Devices and methods for contactless dielectrophoresis for cell or particle manipulation

ABSTRACT

Devices and methods for performing dielectrophoresis are described. The devices contain sample channel which is separated by physical barriers from electrode channels which receive electrodes. The devices and methods may be used for the separation and analysis of particles in solution, including the separation and isolation of cells of a specific type. As the electrodes do not make contact with the sample, electrode fouling is avoided and sample integrity is better maintained.

REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No.12/720,406, filed Mar.9, 2010, which claims priority to U.S. ProvisionalPatent Application 61/158,553, filed Mar.9, 2009, and U.S. ProvisionalPatent Application 61/252,942, filed Oct. 19, 2009, the disclosures ofall of which are hereby incorporated by reference herein.

FIELD OF THE INVENTION

The present invention relates to devices and methods for contactlessdielectrophoresis (DEP) for manipulation of cells or particles. Thedevices and methods of the present invention provide for the applicationof DEP in which electrodes are not in direct contact with the subjectsample.

BACKGROUND OF THE INVENTION

Isolation and enrichment of cells/micro-particles from a biologicalsample is one of the first crucial processes in many biomedical andhomeland security applications [1]. Water quality analysis to detectviable pathogenic bacterium [2-6] and the isolation of rare circulatingtumor cells (CTCs) for early cancer detection [7-19] are importantexamples of the applications of this process. Conventional methods ofcell concentration and separation include centrifugation, filtration,fluorescence activated cell sorting, or optical tweezers. Each of thesetechniques relies on different cell properties for separation and hasintrinsic advantages and disadvantages. For instance, many of the knowntechniques require the labeling or tagging of cells in order to obtainseparation. These more sensitive techniques may require prior knowledgeof cell-specific markers and antibodies to prepare target cells foranalysis.

Dielectrophoresis (DEP) is the motion of a particle in a suspendingmedium due to the presence of a non-uniform electric field [28, 29]. DEPutilizes the electrical properties of the cell/particle for separationand identification [29, 30]. The physical and electrical properties ofthe cell, the conductivity and permittivity of the media, as well as thegradient of the electric field and its applied frequency are substantialparameters determining a cell's DEP response.

The application of dielectrophoresis to separate target cells from asolution has been studied extensively in the last two decades. Examplesof the successful use of dielectrophoresis include the separation ofhuman leukemia cells from red blood cells in an isotonic solution [7],entrapment of human breast cancer cells from blood [8], and separationof U937 human monocytic from peripheral blood mononuclear cells (PBMC)[9]. DEP has also been used to separate neuroblastoma cells from HTBglioma cells [9], isolate cervical carcinoma cells [10], isolate K562human CML cells [11], separate live yeast cells from dead [12], andsegregate different human tumor cells [13]. Unfortunately, themicroelectrode-based devices used in these experiments are susceptibleto electrode fouling and require complicated fabrication procedures [33,34].

Insulator-based dielectrophoresis (iDEP) has also been employed toconcentrate and separate live and dead bacteria for water analysis[2].In this method, electrodes inserted into a microfluidic channel createan electric field which is distorted by the presence of insulatingstructures. The devices can be manufactured using simple fabricationtechniques and can be mass-produced inexpensively through injectionmolding or hot embossing[35, 36]. iDEP provides an excellent solution tothe complex fabrication required by traditional DEP devices however, itis difficult to utilize for biological fluids which are highlyconductive. The challenges that arise include joule heating and bubbleformation[37]. In order to mitigate these effects, oftentimes theelectrodes are placed in large reservoirs at the channel inlet andoutlet. Without an additional channel for the concentrated sample[36],this could re-dilute the sample after it has passed through aconcentration region.

While many have had success designing and fabricating different DEP andiDEP microdevices to manipulate particles in biological fluids, thereare some potential drawbacks of these techniques. The traditional DEPtechnique suffers from fouling, contamination, bubble formation nearintegrated electrodes, low throughput, and an expensive and complicatedfabrication process [33, 34]. The insulating obstacles employed by iDEPare meant to address these shortcomings and are less susceptible tofouling than integrated electrodes [38]. The iDEP fabrication process isalso much less complicated; the insulating obstacles can be patternedwhile etching the microchannel in one step. This technique has the addedbenefit of making the process more economical in that mass fabricationcan be facilitated through the use of injection molding. Unfortunately,one of the primary drawbacks of an iDEP system is the presence of a highelectric field intensity within the highly conductive biological fluidinside the microchannel [33, 39]. The relatively high electrical currentflow in this situation causes joule heating and a dramatic temperatureincrease. The ideal technique would combine the simple fabricationprocess of iDEP and resistance to fouling with the reducedsusceptibility to joule heating of DEP while preserving the cellmanipulation abilities of both methods.

SUMMARY OF THE INVENTION

It is an object of the present invention to provide a dielectrophoresisdevice having a sample channel which is separated by physical barriersfrom electrode channels which receive electrodes. The electrodes providean electric current to the electrode channels, which creates annon-uniform electric field in the sample channel, allowing for theseparation and isolation of particles in the sample. As the electrodesare not in contact with the sample, electrode fouling is avoided andsample integrity is better maintained.

It is a further object of the present invention to provide adielectrophoresis device having a sample channel which is separated byphysical barriers from electrode channels which receive electrodes,whereby the sample channel and electrode channels are formed in a singlesubstrate layer and whereby the physical barriers are formed by thesubstrate itself.

It is a further object of the present invention to provide adielectrophoresis device having a channel for receiving a sample in afirst substrate layer, a first electrode channel and a second electrodechannel for receiving electrodes in a second substrate layer and aninsulation barrier between the first substrate layer and the secondsubstrate layer.

It is a further object of the present invention to provide adielectrophoresis device having a first electrode channel for conductingan electric current in a first substrate layer, a channel for receivinga sample in a second substrate layer and a second electrode channel forconducting an electric current in a third substrate layer. The devicealso has a first insulation barrier between the first substrate layerand the second substrate layer and a second insulation barrier betweenthe second substrate layer and the third substrate layer, preventing thesample from coming in contact with the electrodes.

It is a still further object of the present invention to provide methodsfor separating particles in solution using a device of the presentinvention. A sample containing particles is introduced into the samplechannel in a manner that causes the sample to flow through the channeland electrical current is applied to the electrodes, creating anon-uniform electric field that affects the movement of the particles tobe separated differently than it affects the movement of other particlesin the sample. As the particles to be separated move differently, theyare separated from other particles in the sample at which point they maybe isolated.

There are other objects of the present invention that are provided whichare described in further detail below.

DESCRIPTION OF THE DRAWINGS

FIGS. 1A and B show a three dimensional schematic of a two layer designembodiment of the present invention. The side channels and the mainchannel are fabricated in one layer. FIG. 1B shows and exploded view ofthe box in FIG. 1A.

FIGS. 2A-C show schematics of example electrode channel geometries whichmay be used in embodiments of the present invention. Square (FIG. 2A),rounded (FIG. 2B), and saw-tooth (FIG. 2C) are some examples ofelectrode geometries which may be used in embodiments of the presentinvention.

FIGS. 3A and B show schematics of example embodiments with variations ininsulating barrier geometries in which the barrier thickness (FIG. 3A)increases and decreases (FIG. 3B).

FIGS. 4A-F show schematics of example variations in insulatingstructures within the sample channel which may be used in embodiments ofthe present invention. A single circular structure (FIG. 4A), multipleinsulating structures (FIG. 4B), a diamond shaped insulating structure(FIG. 4C), a ridge insulating structure (FIG. 4D), an oval insulatingstructure (FIG. 4E) and a bump structure (FIG. 4F) are the exampleembodiments shown.

FIGS. 5A-D show schematics of example variations of electrode offsetwhen a single layer device has two electrodes on opposite sides of thesample channel. FIG. 5E is a plot of calculated gradient of electricfield along the center of the sample channel for the various electrodeoffsets.

FIGS. 6A-K show schematics of other embodiments of two layer devicedesigns which may be implemented in embodiments of the presentinvention.

FIGS. 7A-H show schematics of other embodiments of two layer devicedesigns with insulating structures or ridges inside and outside of themain channel.

FIGS. 8A-D show schematics of an embodiment of the three layer device ofthe present invention. FIGS. 8A and B show the layers of the device.FIG. 8C shows a view of the channels taken along section a-a from FIG.8A. FIG. 8D shows an exploded view of the box of FIG. 8B.

FIGS. 9A-D show schematics of an embodiment of the three layer device ofthe present invention. Panels A-D have the same views as are describedfor FIG. 8.

FIGS. 10A-D show schematics of an embodiment of the three layer deviceof the present invention. Panels A-D have the same views as aredescribed for FIG. 8.

FIGS. 11A-C show schematics of an embodiment of the three layer deviceof the present invention for continuous sorting. FIGS. 11A and B showthe layers of the device. FIG. 11C shows a top view of the channels.Tilted electrode channels on the bottom layer are separated from thesample channel with a thin dielectric barrier. The electrodes have anangle with respect to the center line of the main channel. The targetcells can be continuously manipulated in a specific reservoir in theoutlet.

FIG. 12A shows a schematic of an embodiment of a five layer device ofthe present invention. FIG. 12B shows a schematic of a top view of theembodiment of FIG. 12A. FIG. 12C shows a schematic of an embodiment of amultiple layer device of the present invention.

FIG. 13 shows a schematic of an embodiment of a device for continuoussorting having two differently shaped electrodes.

FIG. 14 shows a schematic of an embodiment of continuous sorting devicewith identical electrodes.

FIG. 15 shows a schematic of an embodiment of a batch sorting 5 layerdevice with each electrode and sample channel on a separate layer.

FIG. 16A shows a schematic of an embodiment of a three layer device fortrapping particles. FIG. 16A-1 shows an enlarged view of the dashed boxin FIG. 16A. FIGS. 16B-D show images of red blood cells (FIG. 16B)trapped via positive DEP, 4 micron beads (FIG. 16C) trapped via positiveDEP, and 1 micron beads (FIG. 16D) trapped via negative DEP.

FIGS. 17A-D show schematics of embodiments of three layer devices of thepresent invention. The geometry of the main and side channels may bechanged for different micro-particle DEP manipulation strategies.

FIGS. 18A and B show schematics of an embodiment of a device design tomeasure the electrorotation of the cells/micro-particles suspended inmedium. FIG. 18B shows an exploded view of the box in FIG. 18A.

FIG. 19 shows a circuit diagram of an example electronics system whichmay be used with the devices of the present invention.

FIG. 20 shows a circuit diagram of an example electronics system havinga feedback loop which may be used with the devices of the presentinvention.

FIGS. 21A-F show schematics of a fabrication process which may be usedin conjunction with the present invention. Steps A through D arefollowed only once to create a master stamp. Steps E and F are repeatedto produce an indefinite number of experimental devices. FIG. 21G showsa SEM image of the silicon wafer mold at the intersection between theside and the main channel of the microfluidic device. FIG. 21H shows animaging showing the surface roughness of the wafer after growing andremoving the oxide layer. FIG. 21I shows an image showing the scallopingeffect after DRIE.

FIG. 22A shows a schematic of the microfluidic device of Example 1 andthe equivalent circuit model. FIG. 22A-1 shows an enlarged view of thedashed box in FIG. 22A. FIG. 22B shows a schematic of the two transistorinverter circuit provided by JKL Components Corp.

FIG. 23 shows numerical results of the electric field gradient withinthe sample channel. FIG. 23A shows a surface plot of the gradient of thefield (kg²mC⁻²S⁻⁴) within the main microchannel. FIG. 23B shows a lineplot of the gradient (kg²mC⁻²S⁻⁴) along the line a-b (mm)for fourdifferent frequencies (40, 85, 125, and 200 kHz) at 250 Vrms. FIG. 23Cshows the line plot of the gradient of the electric field along the linea-a for four different applied voltages (100, 200, 350, and 500V) at 85kHz.

FIGS. 24A-C show electric field surface plot for an applied AC field at85 kHz and 250 Vrms. Areas with the induced electric field intensityhigher than (A) 0.1 kV/cm, (B) 0.15 kV/cm and (C) 0.2 kV/cm.

FIGS. 25A and B show superimposed images showing the trajectory of onecell through the device. In FIG. 25A the cell is moving from right toleft under an applied pressure and in FIG. 25B with an applied voltageof 250 Vrms at 85 KHz. The superimposed images were approximately 250 msapart.

FIG. 26 shows a plot of the normalized velocity of THP-1, MCF-7, andMCF-10A cells. U_(on) is the velocity of the cells while applyinge-field and U_(off) is the velocity of the cells while the power is off.

FIG. 27 shows two, single-frame images showing several cells arranged inthe “pearl-chain” phenomena often associated with DEP. These images showthe grouping of cells into a chain configuration in areas of the mainchannel with a high gradient of the electric field. Images were capturedwith an applied field of 250 Vrms at 85 kHz.

FIG. 28A shows a three dimensional schematic of the experimental set upof Example 2. FIG. 28B shows an enlarged view of the center of the threedimensional schematic of FIG. 28A.

FIG. 29A shows two dimensional top view schematic of device 1 of Example2 showing the dominated acting forces on the particle. The contoursrepresent the electric fields modeled in Comsol multiphysics. FIG. 29Bshows a line plot of the gradient of the electric field squared(kg²mC⁻²S⁻⁴) for three different electrical boundary conditions withefficient numerical cell trapping (V1=V2=50 Vrms at 220 kHz, 100 Vrms at152 kHz, and 150 Vrms at 142 kHz and V3=V4=Ground). FIG. 29C shows aline plot of the gradient of the electric field squared (kg²mC⁻²S⁻⁴)along the lines parallel to the center line of the main channel and atdifferent distances from the channel wall for V1=V2=150 Vrms at 140 kHzboundary condition (y=0, 50, and 100 μm).

FIG. 30A shows a two dimensional top view schematic of device 2 ofExample 2, showing the dominated acting forces on the particle. Thecontours represent the electric fields modeled in Comsol multiphysics.FIG. 30B shows a line plot of the gradient of the electric field squared(kg²mC⁻²S⁻⁴)for four different electrical boundary conditions withefficient numerical cell trapping (V1=300 Vrms at 200 kHz, 300 kHz, 400kHz, and 500 kHzV2=Ground) along the x axis (y=0). FIG. 30C shows a lineplot of the gradient of the electric field squared (kg²mC^(Z−2)S⁻⁴) forfour different electrical boundary conditions with efficient numericalcell trapping (V1=30 Vrms at 200 kHz, 300 kHz, 400 kHz, and 500 kHz, andV2=Ground) along the y axis (x=0).

FIGS. 31A-D show plots of: (A) Voltage-frequency pairs to achieve 80%trapping efficiency for device 1 of Example 2; (B) Trapping efficiencyof device2 of Example 2 at 500 kHz and 30 Vrms for flow rates of 0.02,0.04, 0.06, and 0.08 mL/hr; (C) Trapping efficiency at 0.02 mL/hr ofdevice 2 of Example 2 at 200, 300, 400, and 500 kHz as voltages increasefrom 20 Vrms to 50 Vrms; and (D) Maximum gradient of the electric fieldalong the x (y=0) and y (x=0) axis of device 2 of Example 2 forfrequencies between 200 kHz and 1000 kHz.

FIGS. 32A-C show images of experimental results for device 1 of Example2: (A) Dead and live THP-1 cells are moving from right to left due topressure driven flow without applying electric field; (B) 30 secondsafter applying the electric field (V1=V2=100 Vrms at 152 kHz andV3=V4=Ground), the live cells were trapped due to positive DEP, but thedead cells pass by the trapping area; (C) Releasing the trapped livecells by turning off the power supply. Side channels are fluorescent dueto Rhodamine B dye suspended in PBS.

FIGS. 33A-C show images of experimental results for device 2 of Example2: (A) Dead and live THP-1 cells are moving left to right due topressure driven flow; (B) 30 seconds after applying the electric field(V1=40 V_(rms) at 500 kHz and V2=Ground) live cells were trapped due topositive DEP but dead cells pass by; (C) Releasing the trapped livecells by turning off the power supply.

FIGS. 34A-F show a schematic of the fabrication process of Example 3.Steps A through D are followed only once in clean room to create amaster stamp. Steps E and F are repeated to produce an indefinite numberof experimental devices out of clean room and in lab. FIG. 33G shows aSEM image of the silicon wafer mold at the trapping zone. FIG. 33H showsan image of the fabricated device. The main and side channels werefilled with dyes to improve imaging.

FIG. 35A shows an image of a PDMS mold from a silicon master stampcontaining multiple microfluidic devices as described in Example 3. FIG.35B shows a two dimensional schematic of the device with straight mainchannel used in Example 3. The channel depth is 50 μm. FIG. 35C shows anenlarged view of the dashed box in FIG. 35B.

FIG. 36A shows an electric field intensity (V/m) surface plot. FIGS. 36Band C show plots of the gradient of the electric field squared (v(E•E))(kg²mC⁻²S⁻⁴) surface plot. V1=V2=70 Vrms at 300 kHz and V3=V4=Ground.

FIGS. 37A and B show numerical results for Example 3: (A) a line plot ofthe x component of the gradient of the electric field squared(kg²mC⁻²S⁻⁴) along the lines parallel to the center line of the mainchannel and at different distances from the channel wall for V1=V2=70Vrms at 300 kHz and V3=V4=Ground boundary condition (y=0, 50, and 100μm); and (B) a line plot of the y component of the gradient of theelectric field squared (kg2mC-2S-4) along the lines perpendicular to thecenter line of the main channel and at different distances from theorigin for V1=V2=70 Vrms at 300 kHz and V3=V4=Ground boundary condition(x=0, 150, 250, 350, and 450 μm).

FIGS. 38A and B show the gradient of the electric field intensity alongthe centerline of the main channel for different electrodeconfigurations. The electrodes are charged with 70 Vrms and 300 kHz inthe side channels in all cases. Case 1: charged electrodes are inchannels 1 & 2 and ground electrodes are in channels 3 & 4, Case 2:charged electrodes are in channels 1, 2 & 4 and ground electrodes are inchannels 3, Case 3: charged electrodes are in channels 1 & 4 and groundelectrodes are in channels 2 & 3, Case 4: charged electrodes are inchannel 1 and ground electrodes are in channel 2. FIG. 38A shows a plotof the results, while FIG. 38B shows an electric field intensity surfaceplot.

FIG. 39 shows an image of experimental results from Example 3, a brightfield image of live THP-1 cells, shown here 30 seconds after applyingthe electric field (V1=V2=70 Vrms at 300 kHz and V3=V4=Ground). Thecells were trapped due to positive DEP.

FIGS. 40A-C show images of experimental results from Example 3:selective trapping of live THP-1 cells from a mixture also containing 10μm polystyrene beads. THP-1 live cells were stained using cell tracecalcein red-orange dye (A) Cells and beads are moving from right to leftdue to pressure driven flow. (B) THP-1 cells are trapped viadielectrophoresis and beads are passing through the trapping zone.Charged electrodes are in channels 1 & 2 (V1=V2=70 Vrms) at 300 kHz andground electrodes are in channels 3 & 4 (V3=V4=G). (C) Releasing thetrapped cells.

FIGS. 41A-D show images of experimental results from Example 3: trapping2 μm beads suspended in DI water (V1=V2=190 Vrms at 300 kHz andV3=V4=Ground) (A) t=0 (B) t=30 Seconds (C) t=50 Seconds (D) t=1 min,Release.

FIGS. 42A and B show schematics of a device designed for continuousseparation of particles as is described in Example 4, with FIG. 42Bshowing an exploded view of the box in FIG. 42A. Particles are driventhrough the sample channel while an electric signal is applied acrossthe fluid electrode channels. Four micron beads are continuouslyseparated from 2 micron beads and released, as is shown in the images ofFIGS. 42C and D. Red blood cells are isolated from buffer solution, asis shown in the image of FIG. 42E.

DETAILED DESCRIPTION OF THE INVENTION

The present invention provides methods, devices, and systems tomanipulate micro-particles suspended in biological fluids using theirelectrical signatures without direct contact between the electrodes andthe sample. Contactless dielectrophoresis (cDEP) employs the simplifiedfabrication processes of iDEP yet lacks the problems associated with theelectrode-sample contact [40].

cDEP relies upon reservoirs filled with highly conductive fluid to actas electrodes and provide the necessary electric field. These reservoirsare placed adjacent to the main microfluidic channel and are separatedfrom the sample by a thin barrier of a dielectric material. Theapplication of a high-frequency electric field to the electrodereservoirs causes their capacitive coupling to the main channel and anelectric field is induced across the sample fluid.

Similar to traditional DEP, cDEP may exploit the varying geometry of theelectrodes to create spatial non-uniformities in the electric field.However, by utilizing reservoirs filled with a highly conductivesolution, rather than a separate thin film array, the electrodestructures employed by cDEP can be fabricated in the same step as therest of the device; hence the process is conducive to mass production[40]. The various embodiments of the present invention provide devicesand methods for performing cDEP, as well as methods for fabricating cDEPdevices.

In general, the present invention provides devices and methods thatallow cell sorting to identify, isolate or otherwise enrich cells ofinterest based on electrical and physical properties. An electric fieldis induced in a main sorting microchannel using electrodes inserted in ahighly conductive solution which is isolated from the microchannel bythin insulating barriers. The insulating barriers exhibit a capacitivebehavior and an electric field is produced in the isolated microchannelby applying an AC electric field. Electrodes do not come into contactwith the sample fluid inside the microchannel, so that electrolysis,bubble formation, fouling and contamination is reduced or eliminated. Inaddition, the electric field is focused in a confined region and has amuch lower intensity than that found in traditional insulator-baseddielectrophoresis, so heating within the sample channel is negligibleand the likelihood of cell lysis is greatly reduced. The system can alsobe used for characterizing and sorting micro-or nanoparticles.

Methods

In one embodiment, the present invention provides a method to induce DEPto manipulate cells or micro/nano particles without direct physicalcontact between the electrodes and the sample solution with a simplifiedand inexpensive micro-fabrication process. Further examples ofmanipulation of cells and micro/nano particles are given below.

In another embodiment, the present invention provides a method to inducean electric ac field without direct physical contact between theelectrodes and the sample solution with a simplified and inexpensivemicro-fabrication process.

In another embodiment, the present invention provides a method wherebycDEP can be used to measure the current through a system and measure theelectrical resistance/impedance of a system for detection purposes. cDEPelectrodes can be placed on an object to deliver a known amount ofelectrical current though the object. By measuring the electricpotential at different places on the object, the electrical impedance ofthe object can be calculated. In this embodiment, the electricalimpedance may be measured so that it is possible to determine when acertain number of particles are trapped or isolated. Once the requisitenumber of particles are trapped, e.g. the number required for downstreamanalysis, the impedance will reach a pre-set level and the current canbe turned off, allowing the particles to be released.

In another embodiment of the present invention, cDEP can be used as anon-invasive method to monitor living animal cells in vitro. The cellsare grown on an insulating thin barrier. The electrode channels areunder this thin barrier. The impedance of the cultured cells on theinsulating barrier is measured at one specific frequency as a functionof time. Because of the insulating properties of the cell membrane, theimpedance of the system increases with increasing the number of cells onthe surface. The 3D geometrical changes of layered cells on the surfacecan be monitored because the current through the layers of cells andaround the cells changes due to the shape change of the cells.

In another embodiment of the present invention, methods are providedwhereby cDEP can be used to measure the dielectric properties of amedium as a function of frequency. The impedance of a electrochemicalsystem is measured for different frequencies to characterize theresponse of the system as a function of frequency

In another embodiment of the present invention, cDEP devices can bedesigned to provide methods for measuring small changes in electricalresistance of the chest, calf or other regions of the body withoutdirect electrode-body contact to monitor blood volume changes. Thesemethods can indirectly indicate the presence or absence of venousthrombosis and provide an alternative to venography, which is invasiveand requires a great deal of skill to execute adequately and interpretaccurately.

In yet another embodiment of the present invention, cDEP devices may beused for solution exchange and purification of particles. As anon-limiting example, once the particles of interest are captured in adevice, the inlet solution may be change to a solution different fromthat of the sample, for example a buffer. The particles may be releasedinto the buffer. As a non-limiting example, cancer cells may beconcentrated from a blood sample in the device. The inlet solution maythen be changed to a suitable buffer, allowing the cancer cells to bepurified and concentrated from blood and suspended in the buffer.

In still another embodiment of the present invention, cDEP devices maybe designed to have two (or more) solutions traveling side by side usinglaminar flow as is known in the art. Changes in the electrical field ofthe device may then be used to move particles back and forth between thetwo flows as is necessary. The two flows may then later be separated sothat particles are isolated as desired.

The methods of the present invention may involve any DEP deviceengineered so that there is no direct physical contact between theelectrodes and the sample solution. Exemplary, but non-limiting,examples of such devices are given in this specification.

Device Designs

Non-limiting examples of cDEP device designs are presented herein. Someexamples are illustrated in the figures, where like numbering may beused to refer to like elements in different figures (e.g. element 117 inFIG. 1 may have a similar function to element 217 in FIG. 2). Theobjects and elements shown in a single figure may or may not all bepresent in one device. The present invention contemplates any DEP deviceengineered so that there is no direct physical contact between theelectrodes and the sample solution, and there will be modifications ofthe examples set forth herein that will be apparent to one of skill inthe art.

One Layer (2D) Designs

In certain embodiments of the invention, a device is provided where themain and side (electrode) channels are fabricated in one layer of thedevice. The second layer is an insulating layer such as glass orpolydimethylsiloxane (PDMS) to bond the microfluidic channels.

FIG. 1A shows a 3D schematic example of a 2D cDEP device 111 with themain and side channels fabricated in one layer. Side channel electrodes113, 115 and the main sample channel 117 are fabricated in a singlesubstrate layer 119. FIG. 1B shows an exploded view of the box in FIG.1A, where notches 121 in the electrode channels 113, 115 can be seen.The electrode channels have portions of receiving electrodes 114, 116,which are shown as circular but may be different shapes depending on theelectrode to be received. It is further contemplated that the electrodechannels need not have specially shaped portions for receiving anelectrode, as the electrode can simply be contacted with the conductivesolution in the channel.

There are many factors affecting the performance of single-andmulti-layer devices. These factors include the electrode channelgeometry, insulating barrier thickness, insulating barrier geometry,insulating structures within the sample channel, sample channel width,sample channel depth, distance between electrodes, and number ofelectrodes. These factors may be modified to customize the electricfields inducing DEP.

The electrode channels may have a variety of shapes and sizes whichenhance the performance of single- and multi-layer devices. Exampleshapes include: square or rectangular electrodes, rounded squares orrectangles (radius of curve additionally effects performance), saw-toothshapes, combinations of these shapes or any geometric change to theelectrode channel. For the purposes of the invention, symmetry is notrequired and asymmetry can alter the performance of the device. Examplesof rectangular electrodes 223 (FIG. 2A), rounded rectangular electrodes225 (FIG. 2B) and saw tooth shaped electrodes 227 (FIG. 2C) on eitherside of sample channels 217 are shown in FIG. 2. It should be apparentthat other rectangular, rounded rectangular and saw-tooth shapedelectrodes are contemplated by the present invention and that theembodiments in FIG. 2 are exemplary only.

Insulating barrier thickness is the thickness of the insulating materialwhich separates the electrode channels and the sample channel. Thethickness of the insulating barrier can change the performance of thedevice. In certain embodiments, these thicknesses can vary between about0.01 micron and about 10 mm, and are preferably between about 1 micronand about 1000 micron. It is contemplated that each electrode channelmay have a different insulating barrier thickness.

The geometry of the insulating barriers may change the performance ofthe device. Some contemplated variations include: straight barriers,increases or decreases in barrier thickness along the length (FIG. 3),rounded barriers, barriers which become thicker or thinner along thedepth of the channel and combinations of these variations. As is shownin FIG. 3, certain embodiments of devices of the present invention mayhave areas where the thickness of the insulation barrier increases 329(FIG. 3A) or where the thickness of the insulation barrier decreases 331(FIG. 3B).

It is further contemplated that insulating structures may be present inthe sample channel or the electrode channels to affect the electricalfield. The insulating structures may consist of many different shapesand sizes, including: round or cylindrical pillars, ridges or shelveswhich split the channel, bumps or slope changes along the channel wallsor floors and other geometric changes within the channel (see FIGS. 4and 7).

FIG. 4 shows non-limiting examples of insulating structures which may beused in the devices of the present invention: a single round post 433(FIG. 4A), double round posts 433 (FIG. 4B), square posts 435 (FIG. 4C),angled shapes 437 (FIG. 4D), rounded rectangles 439 (FIG. 4E) andextensions of the insulating barrier into the sample channel 441 (FIG.4F). It will be apparent to one of skill in the art that there areextensive variations on the embodiments shown in FIG. 4 that fall withinthe scope of the present invention.

The sample channel width may change the performance of the device. Incertain embodiments, this width may vary between about 1 micron andabout 10 cm, and is preferably between about 10 micron and about 1000micron.

The sample channel depth may also change the performance of the device.In certain embodiments, this depth may vary between about 1 micron andabout 10 cm, and is preferably between about 10 micron and about 1000micron.

Electrode offset, or the distance between electrodes is another designfactor which may change the performance of the device. In certainembodiments, this offset may vary between no offset and about 10 cmoffset, but is ideally between 0 micron and about 1 mm. The effects ofthis offset can be seen in FIG. 5 which shows electrode offsets of 0micron (FIG. 5A), 50 micron (FIG. 5B), 100 micron (FIG. 5C) and 200micron (FIG. 5D). As is shown in the plot of FIG. 5E, the calculatedgradient of electric field along the center of the sample channelincreases as the offset is increased from 0 microns to 200 microns.Above this offset, the electric field gradient decreases. It should benoted that this behavior is for the design with a 100 micron sample, 20micron barriers, and 100 micron wide electrode channels. As will beapparent to one of skill in the art, different geometries will havedifferent responses to offsets.

It will be apparent to one of skill in the art that many other cDEPdevices with different geometries and strategies to manipulatemicro-particles fall within the scope of the present invention.Additional, non-limiting embodiments of 2D devices of the presentinvention are shown in FIG. 6A-K, with sample channels 617, electrodes613, 615 and insulating structures 643 as illustrated.

The insulating structures and ridges inside and outside of the mainchannel can be used to enhance the cDEP effect. cDEP separation ofmicro/nano-particles strongly depends on the geometry of thesestructures. In certain embodiments, insulating structures within thesample channel may consist of many different shapes and sizes,including: round or cylindrical pillars, ridges or shelves which splitthe channel, bumps or slope changes along the channel walls or floorsand other geometric changes within the sample channel. It is alsocontemplated that on or both of the electrode channels may haveinsulating structures.

Non-limiting examples of different cDEP devices showing differentstrategies to use these insulating structures inside and outside of themain channels are shown in FIGS. 7 A-H, with sample channels 717,electrodes 713, 715 and insulating structures 743 as illustrated. FIG.7C shows an embodiment with insulating structures in the sample channeland the electrode channels.

Three Layer Designs

In other embodiments of the invention, the main channel and theelectrode channels are fabricated in two separate insulating layers. Thethird layer is a thin insulating barrier separating the other twolayers. In certain embodiments, the insulating barrier is made frompoly(methyl methacrylate) (PMMA). In other embodiments of the invention,the insulating barrier is made from plastic, silicon, glass,polycarbonate, or polyimide, such as the polyimide film KAPTON producedby Dow Chemical (Midland Mich.). Specific, non-limiting examples includesilicon oxide, silicon nitride and polyethylene. The geometry, shape,and position of the bottom or top electrode channels are importantparameters in cell/microparticle manipulation. Four non-limitingexamples of such designs are shown in FIGS. 8-11.

FIGS. 8-10 show embodiments of three layer devices of the presentinvention, with panels A and B of each figure showing view of the layersof the device, panel C showing a view of the overlap of the channelsalong section a-a of panel A and panel D showing an exploded view of theboxed area of panel B. As is shown in FIG. 8, the sample channel isfabricated in the sample channel layer 845, while the electrode channels813, 815 are fabricated in the electrode channel layer 849. Theinsulating barrier 847 separates the sample channel layer 845 andelectrode channel layer 849. As is shown in the embodiment of FIG. 8,the sample channel layer 845 has holes for accessing the sample channel846 as well as holes for receiving electrodes 848, 850. Holes forreceiving electrodes 848, 850 are also present in the insulating barrier847 so that the electrodes may make contact with the electrode receivingportions 814, 816 of the electrode channels 813, 815. FIGS. 9 and 10show other embodiments with like numbering representing like elements.

FIGS. 11A-C show a schematic of a three layer device with electrodechannels 1113, 1115 on the bottom substrate layer 1149 and the samplechannel 1117 on the top substrate layer 1145. A syringe pump 1152 isused in the embodiment of FIG. 11 for injecting the sample into thesample channel. The electrode channels 1113, 1115 and the sample channel1117 are separated with a thin insulating barrier 1147. The anglebetween the electrode channels 1113, 1115 and the sample channel 1117can be adjusted between 0 to 90 degree. These tilted electrode channels1113, 1115 manipulate the cells or micro-particles towards the sides ofthe sample channel 1117 such that the target particles along the side ofthe sample channel 1117 can be collected in a separate reservoir. As isshown in FIG. 11A, target particles may be separated and isolated in atarget reservoir 1156, while the remaining particles in the sample flowinto the normal reservoir 1158. FIG. 11C shows the top view of just thesample channel 1117 and the electrode channels 1113, 1115 for the deviceshown in FIG. 11A.

Five Layer Designs

In other embodiments of the invention, a five layer device may be used.These designs have a sample channel with electrodes above and below it.A thin membrane above and below the sample channel isolate it from theelectrode channels. An non-limiting example of this embodiment can beseen in FIG. 12A. The embodiment shown in FIG. 12A has a top cover 1251,an electrode channel layer 1249 with an electrode channel 1213, aninsulating layer 1247, a sample channel layer 1245 with a sample channel1217, an insulating layer 1247, an electrode channel layer 1249 with anelectrode channel 1215 and a bottom cover 1253. FIG. 12B shows aschematic representing a top view of the device shown in FIG. 12A, withthe overlapping electrode channels 1213, 1215 and the sample channel1217 shown.

Multiple Layers

In other embodiments of the present invention there are providedmultiple layer cDEP devices. These designs consist of multiple samplechannels within one device. They may be organized in layers as:electrode—barrier—sample—barrier—electrode—barrier—sample—barrier, withthe pattern repeating. Those skilled in the art of fabrication will beable to create devices with upward of 10 sample channels in a singledevice. An example of this configuration with three sample channels canbe seen in FIG. 12C. FIG. 12C shows alternating electrode layers 1249containing an electrode 1213 or 1215, insulating layers 1247, and samplechannel layers 1245 containing a sample channel 1217. The layers aresandwiched between a top cover 1251 and a bottom cover 1253.

Other Embodiments

The embodiment depicted in FIG. 13 is a three layer device. Bothelectrodes 1313, 1315 are located in the same layer. They are separatedfrom the sample channel 1317 by an insulating layer 1347. The entiredevice is encased within a non-conducting case, which is not shown. Inthis device, particles traveling in the straight part of the samplechannel (the part of the sample channel parallel to the gap betweenelectrodes) will be diverted by dielectrophoretic forces. Particles withspecific electrical properties will be diverted into the T-section ofthe sample channel (the part of the sample channel perpendicular to thegap between electrodes) while others continue straight. Devices of thisnature will continuously sort particles as they flow through the device.

The embodiment depicted in FIG. 14 is a three layer device with bothelectrodes 1413, 1415 located on the same layer. In the embodiment ofFIG. 14, the sample channel 1417 splits into two channels, and uppersample channel 1455 and a lower sample channel 1457. In this embodiment,as they travel from right to left, particles experiencing positive DEPwill be deviated into the upper sample channel 1455 while particlesexperiencing negative DEP will be forced into the lower channel 1457,allowing for their separation.

The boundary and material properties depicted are typical of thosetested experimentally.

The embodiment depicted in FIG. 15 is a five layer device. A thinmembrane separates the bottom electrode 1513 from the sample channel1517 and another separates the sample channel 1517 from the topelectrode 1515. This device may be used to batch sort particles. An ACelectric field is applied to the electrodes. Particles would be allowedto trap in the region where the electrodes overlap 1559. After a desiredtime, the electric field would be reduced releasing the particles fordownstream analysis.

The embodiment depicted in FIG. 16 is a three layer device. A 50 micronPMMA barrier separates the electrode channels 1613, 1615 from the samplechannel 1617. In this embodiment, the two electrode channels areseparated by 100 microns and each channel is 500 microns wide. Thisdesign can be used to batch sort cells and continuously sort cells.Below a certain threshold, particles may be pushed toward one side ofthe sample channel, separating them from the bulk solution (continuoussorting). Above a certain threshold, particles will be trapped in theregion of the sample channel which lies between the two electrodechannels. FIG. 16B shows an image of pearl chaining red blood cellsbeing trapped in the sample channel at 200 kHz and 50 V. FIGS. 16C and Dshow images of 4 micron beads being trapped along the sample channelwalls while 1 micron beads are forced to the center of the channel bynegative DEP at 400 kHz and 50 V.

Further, non-limiting examples of embodiments of devices of the presentinvention are shown in FIGS. 17A-D, with like numbers indicating likeelements.

In certain embodiments of the present invention, the sample channel maybe designed with multiple inlets and outlets. Multiple inlets andoutlets for the sample channel may allow the cDEP device moreflexibility for sample handling and micro-particle manipulation fordifferent purposes.

The methods and devices of the present invention allow for the sort ofvarious types of particles, including cells. For the purposes of thisdisclosure sorting is intended to mean the separation of particles basedon one or more specific characteristics. There are many differentcharacteristics by which particles may be sorted, including, but notlimited to: particle size, particle shape, particle charge, internalconductivity, shell or outer layer conductivity, proteins present in oron the particle, genetic expression, ion concentrations within theparticle, state—for example metastatic vs non-metastatic cancer cells ofthe same phenotype and cellular genotype. Particles that may beseparated, isolated and/or analyzed using the methods and devices of thepresent invention include cells isolated from organisms, single celledorganisms, beads, nanotubes, DNA, molecules, few cell organisms(placozoans), Zygotes or embryos, drug molecules, amino acids, polymers,monomers, dimers, vesicles, organelles and cellular debris.

The methods by which certain embodiments sort particles can vary butinclude: batch sorting (where particles of a certain type are trapped ina particular region for a time before being released for lateranalysis), continuous sorting (where particles of a certain type arecontinuously diverted into a separate region of the channel or device),repulsion (negative DEP), attraction (positive DEP), and field flowfractionation.

cDEP and Downstream Analysis

cDEP can be used in combination with other microfluidic technologies toform complete lab on a chip solutions. Examples of some downstreamanalysis techniques include: flow cytometry, PCR and impedancemeasurement, which may be used alone or in combination. Those of skillin the art will recognize that there are other methods of downstreamanalysis that may be applied after particles are sorted using thedevices and methods of the present invention.

The devices and methods of the present invention can be used to enhanceother trapping and sorting technologies such as dielectrophoresis,insulator based dielectrophoresis (iDEP), protein marker detection,field flow fractionation and diffusion (e.g. H-channel devices). Forexample, a device may have insulating pillars coated with a particularbinding protein to detect circulating cancer cells. However, it isnecessary that cells come in contact with the pillars in order for themto become permanently attached. cDEP can be employed to ensure thatparticles come in contact with the pillars, thus trapping anycirculating cancer cells even after the electric field is removed.

Conductive Solutions

Any conductive solution or polymer may be used in the electrode channelsof devices of the present invention. Examples of conductive solutionsinclude phosphate buffer saline (PBS), conducting solutions, conductivegels, nanowires, conductive paint, polyelectrolytes, conductive ink,conductive epoxies, conductive glues and the like.

Fluid Flow

In certain embodiments, pressure driven flow or electrokonetic flow canbe used to move the sample in the sample channel. The pressure drivenflow used may be provided by an external source, such as a pump orsyringe, or may be provided by the force of gravity. One of skill in theart will recognize that various methods are applicable for moving thesample in the sample channel.

Electrorotation Rate Measurement (ROT Spectra)

It is contemplated that cDEP devices may be designed to measure theelectrorotation rate of different cell lines/micro-particles atdifferent frequencies. These measurements can be used to back out theelectrical properties of the cells/micro-particles. Methods formeasuring such rates will be known to one of skill in the art. FIGS. 18Aand B show an embodiment of the present invention which may be used formeasurement of electrorotation rate, with panel B showing an explodedview of the region in the box in panel A. The sample channel 1817 issurrounded by pairs of each electrode 1813, 1815.

Electrorotation relies on a rotating electric field to rotate the cellsor micro-particles. The electrical properties of the cells ormicro-particles can be calculated by measuring the rotation speed of theparticles at different applied frequencies. The rotating field isproduced by electrodes arranged in quadrupole as shown in FIG. 18B. Theelectrodes are energized with AC signals phased 0°, 90°, 180°, and 270°.

cDEP and Electroporation

Reversible electroporation is a method to temporarily increase the cellmembrane permeability via short and intense electrical pulses. Thedevices of the present invention may be designed to immobilize targetcells in a medium dielectrophoretically with minimum mechanical stresseson the cell and reversibly electroporate the trapped cell. Theconductivity of the cell is changed after electroporation. The devicecan be designed such that the electroporated cell leaves the trappingzone.

Irreversible electroporation (IRE) is a method to permanently open upelectropores on the cell membrane via strong enough electrical pulses.The devices of the present invention may be designed to trap targetcells using dielectrophoresis at trapping zones. These devices may bedesigned such that there is strong enough electric field at the trappingzone to irreversibly electroporate the trapped cell. The conductivity ofthe dead cell changes dramatically and therefore the DEP force decreasesand the target cell can be released after IRE.

Electronics Used with Contactless Dielectrophoresis

In certain embodiments of the present invention, a sinusoidal signal maybe used to elicit a DEP response from particles in the device. However,any electrical signal or signals that capitalize upon the capacitivenature of the barriers between the electrodes and fluidic channel(s) maybe used with the present invention. These include sinusoidal, square,ramp, and triangle waves consisting of single or multiple fundamentalfrequencies however those familiar with electrical signal generationwill be able to develop time-varying signals that may be used. Thefrequency range used to induce a DEP response in may range from tens ofkilohertz to the megahertz range. However, it is also contemplated thatdevices may be designed to utilize frequencies range of several hundredhertz to hundreds of megahertz. For some of the embodiments presentedherein, signal amplitudes ranged from about 30V(peak) to about500V(peak). The amplitude of the applied signal only needs to be of amagnitude that induces a sufficient electric field in the channel tocause a change in cell behavior. Thus the required amplitude of thesignal is dependent on the device configuration and DEP response of thetarget (cell, micro-particle, etc.).

There are numerous methods to generate a signal that may be used forcontactless dielectrophoretic manipulation of cells and micro-particles.Methods for signal generation include oscillators (both fixed andvariable), resonant circuits, or specialized waveform generationtechnologies including function generators, direct digital synthesisICs, or waveform generation ICs. The output of these technologies may becomputer controlled, user controller, or self-reliant.

The output of a signal generation stage may then be coupled to thecontactless dielectrophoretic device directly or coupled with anamplification technique in order to achieve the necessary parameters(voltage, current) for use in a device. Methods for amplificationinclude solid state amplifiers, integrated circuit-based amplifiers,vacuum tube-based technologies, and transformers. Also, diode-basedswitches, semi-conductor devices used in the switch-mode, avalanchemode, and passive resonant components configured to compress and/oramplify a signal or pulse may be used to create a signal(s) to be usedin contactless dielectrophoretic devices. An example electronics systemwhich may be used with the devices of the present invention is shown inFIG. 19. In this implementation, a common laboratory function generatoris used to generate the time varying signal necessary forexperimentation. This signal is input to a solid-state amplifier whichperforms preliminary voltage and current amplification. Further voltageamplication is provided by inputting the output of the amplifier into ahigh voltage transformer which is then coupled to the electrode channelsof the device.

Signal generation technology implemented with a feedback control systemwhich allows the direct control of the electric field parameters withinthe device (electric field intensity, phase, frequency). One possibletopology of a feedback implementation which may be used with the presentinvention is shown in FIG. 20. In FIG. 20 the current passing throughthe cDEP device is being measured in order to determine the magnitude ofthe electric field present within the device. There are several methodsto perform this measurement including, but not limited to, current shuntresistors, current transformers, and transimpedance amplifiers. Themeasured current through the cDEP device is then used to maintain a theelectric field in the device by adjusting the level of the signalgeneration or the gain of the amplification stages. However, thoseproficient in electrical engineering will be able to develop otherfeedback loop implementations to control the parameters of the electricfield within the device.

The devices of the present invention may be coupled with othertechnologies to expand the functionality of the system. This may includeadditional electronics such as rotational spectroscopy or impedancedetection in order to produce systems with a wider range offunctionality.

Fabrication of Devices

The devices of the present invention may be fabricated using astamp-and-mold method. An exemplary illustrated process flow is shown inFIG. 21. A silicon wafer is patterned using photolithography (FIGS. 21Aand B) and then etched using deep reactive ion etching (DRIE) (FIGS. 21Cand D). This etched wafer then serves as a mold onto whichpolydimethylsiloxane (PDMS) is poured and then allowed to cure FIG.21E). The cured PDMS is then removed from the silicon wafer and containsan imprint of the device. Fluid ports are then punched in the cured PDMSmold as needed. Finally, the PDMS mold of the device is bonded to aglass microscope slide using oxygen plasma (FIG. 21F) and fluidicconnections are punched through the PDMS.

Those skilled in microfabrication techniques will be able to modify thisfabrication process to take advantage of materials with propertiesadvantageous to the devices of the present invention. For example, themicrofluidic structures of the device may be etched into a wafer ofdoped or intrinsic silicon, glass (such as Pyrex), or into an oxidationor nitride layer formed on top of a wafer. These materials would allow aresearcher to perform experiments over a wider range of voltages andfrequencies due to their increased permittivity and dielectric strength.Furthermore, the devices of the present invention lend themselves toother production techniques more suitable for mass fabrication such asinjection molding and hot embossing.

It is also contemplated that there are other embodiments such asmicromachining and capillary effect with glass beads that are notexplicitly shown but that someone familiar with the art may employ inpracticing the present invention.

Further specific examples of embodiments of the present invention areshown below. These examples are provided for exemplary purposes only andshould not be considered to limit the scope of the invention as is setforth in the claims below.

EXAMPLES Example 1 Separation of Cells Using cDEP

Background

Efficient biological particle separation and manipulation is a crucialissue in the development of integrated microfluidic systems. Currentenrichment techniques for sample preparation include density gradientbased centrifugation or membrane filtration (57), fluorescent andmagnetic activated cell sorting (F/MACS) (61), cell surface markers(55), and laser tweezers (49). Each of these techniques relies ondifferent cell properties for separation and has intrinsic advantagesand disadvantages. Typically more sensitive techniques may require priorknowledge of cell-specific markers and antibodies to prepare targetcells for analysis.

One alternative to these methods is dielectrophoresis (DEP) which is themotion of a particle due to its polarization in the presence of anon-uniform electric field (28,29). Currently, typical dielectrophoreticdevices employ an array of thin-film interdigitated electrodes placedwithin the flow of a channel to generate a non-uniform electric fieldthat interacts with particles near the surface of the electrode array(63). Such platforms have shown that DEP is an effective means toconcentrate and differentiate cells rapidly and reversibly based ontheir size, shape, and intrinsic electrical properties such asconductivity and polarizability. These intrinsic properties arise due tothe membrane compositional and electrostatic characteristics, internalcellular structure, and the type of nucleus (56) associated with eachtype of cell.

The application of dielectrophoresis to separate target cells from asolution has been studied extensively in the last two decades. Examplesof the successful use of dielectrophoresis include the separation ofhuman leukemia cells from red blood cells in an isotonic solution (7),entrapment of human breast cancer cells from blood (8), and separationof U937 human monocytic from peripheral blood mononuclear cells (PBMC)(9). DEP has also been used to separate neuroblastoma cells from HTBglioma cells (9), isolate cervical carcinoma cells (10), isolate K562human CML cells (11), separate live yeast cells from dead (12), andsegregate different human tumor cells (13). Unfortunately, themicroelectrode-based devices used in these experiments are susceptibleto electrode fouling and require complicated fabrication procedures(33,34).

Insulator-based dielectrophoresis (iDEP) is a practical method to obtainthe selectivity of dielectrophoresis while overcoming the robustnessissues associated with traditional dielectrophoresis platforms. iDEPrelies on insulating obstacles rather than the geometry of theelectrodes to produce spatial non-uniformities in the electric field.The basic concept of the iDEP technique was first presented by Masuda etal. (60). Others have previously demonstrated with glass insulatingstructures and AC electric fields that iDEP can separate DNA molecules,bacteria, and hematapoietic cells (64). It has been shown thatpolymer-based iDEP devices are effective for selective trapping of arange of biological particles in an aqueous sample (51). The patternedelectrodes at the bottom of the channel in DEP create the gradient ofthe electric field near the electrodes such that the cells close enoughto the bottom of the channel can be manipulated. However, the insulatorstructures in iDEP that usually transverse the entire depth of thechannel provide non uniform electric field over the entire depth of thechannel. iDEP technology has also shown the potential for water qualitymonitoring (35), separating and concentrating prokaryotic cells andviruses (58), concentration and separation of live and dead bacteria(2), sample concentration followed by impedance detection (36), andmanipulation of protein particles (59).

While many have had success designing and fabricating different DEP andiDEP microdevices to manipulate particles in biological fluids, thereare some potential drawbacks of these techniques. The traditional DEPtechnique suffers from fouling, contamination, bubble formation nearintegrated electrodes, low throughput, and an expensive and complicatedfabrication process (33,34). The insulating obstacles employed by iDEPare meant to address these shortcomings and are less susceptible tofouling than integrated electrodes (38). iDEP's fabrication process isalso much less complicated; the insulating obstacles can be patternedwhile etching the microchannel in one step. This technique has the addedbenefit of making the process more economical in that mass fabricationcan be facilitated through the use of injection molding.

Unfortunately, one of the primary drawbacks of an iDEP system is thepresence of a high electric field intensity within the highly conductivebiological fluid inside the microchannel (33, 39). The relatively highelectrical current flow in this situation causes joule heating and adramatic temperature increase. The ideal technique would combine iDEP'ssimple fabrication process and resistance to fouling with DEP's reducedsusceptibility to joule heating all-the-while preserving the cellmanipulation abilities of both methods.

The inventors have developed an alternative method to provide thespatially non-uniform electric field required for DEP in whichelectrodes are not in direct contact with the biological sample. Theabsence of contact between electrodes and the sample fluid inside thechannel prevents bubble formation and mitigates fouling. It is alsoimportant to note that without direct contact between the electrodes andthe sample fluid, any contaminating effects of this interaction can beavoided. In fact, the only material in contact with the sample fluid isthe substrate material the device is patterned on. In the presentmethod, an electric field is created in the microchannel usingelectrodes inserted in a highly conductive solution which is isolatedfrom the main channel by thin insulating barriers. These insulatingbarriers exhibit a capacitive behavior and therefore an electric fieldcan be produced in the main channel by applying an AC electric fieldacross them. Furthermore, non-uniformity of the electric fielddistribution inside the main channel is provided by the geometry ofinsulating structures both outside and inside the channel.

In order to demonstrate this new method for cell separation andmanipulation, a microfluidic device to observe the DEP response of cellsto a non-uniform electric field created without direct contact fromelectrodes has been designed and fabricated. Modeling of the non-uniformelectric field distribution in the device was accomplished through anequivalent electronic circuit and finite element analysis of themicrofluidic device. The effects of different parameters such as totalapplied voltage, applied frequency, and the electrical conductivity ofthe fluid inside and outside of the main channel on the resulting DEPresponse were simulated and then observed through experimentation. A DEPresponse was observed primarily as a change in cell trajectory orvelocity as it traveled through the device. Further evidence of this DEPresponse to the non-uniform electric field is provided by theelectrorotation of cells, and their aggregation in “pearl chain”formations.

Theory

Dielectrophoresis DEP is the motion of polarized particles in a nonuniform electric field toward the high (positive DEP) or low (negativeDEP) electric field depending on particle polarizability compared withmedium conductivity. The time-average dielectrophoretic force isdescribed as (28,29):

F _(DEP)=2πε_(m) r ³Re{K(ω)}∇(E _(rms) ·E _(rms))   (1)

where ε_(m) is the permittivity of the suspending medium, r is theradius of the particle, E_(rms) is the root mean square electric field.Re{K(ω)} is the real part of the Clausius-Mossotti factor K(ω) . TheClausius-Mossotti is given by:

$\begin{matrix}{{K(\omega)} = \frac{ɛ_{p}^{*} - ɛ_{m}^{*}}{{ɛ_{p}^{*} + {2ɛ_{m}^{*}}}\;}} & (2)\end{matrix}$

where ε*_(p) and ε*_(m) are the complex permittivities of the particleand the medium, respectively. Complex permittivity is defined as

$\begin{matrix}{ɛ^{*} = {ɛ + \frac{\sigma}{j\omega}}} & (3)\end{matrix}$

where ε, and σ are the real permittivity and conductivity of the subjectand ω is the frequency.

Electrorotation is the rotation of polarized particles suspended in aliquid due to an induced torque in a rotating electric field (37). Themaximum magnitude of the torque is given by

Γ=−4πε_(m) r ³Im{K(ω)}(E _(rms) ·E _(rms))   (4)

where Im{K(ω)} is the imaginary part of the Clausius-Mossotti factorK(ω).

Assuming the cells are spherical particles in the medium, thehydrodynamic frictional force, f_(Drag), due to translation andhydrodynamic frictional torque, R, due to rotation are given by:

f _(Drag)=6ηrπ(u _(p) −u _(f))   (5)

R=8ηr³πχ  (6)

where r is the particle radius, η is the medium viscosity, u_(p) is thevelocity of the particle, u_(f) is the medium velocity, R is inducedtorque, and Ω is electrorotation rate (rad.S⁻¹).

The magnitude of the steady state electrorotation rate Ω andtranslational velocity is determined by a balance between the inducedtorque and the hydrodynamic friction and between the induceddielectrophoretic force and Stoke's drag force on a cell respectively.In this preliminary study it should be noted that the effect of theacceleration term is considered to be negligible. The relationship isgiven by:

$\begin{matrix}{{\Omega (\omega)} = {\frac{ɛ_{m}}{2\eta}{{Im}\left( \frac{ɛ_{p}^{*} - ɛ_{m}^{*}}{{ɛ_{p}^{*} + {2ɛ_{m}^{*}}}\;} \right)}{E_{rms} \cdot E_{rms}}}} & (7) \\{u_{p} = {u_{f} - {\mu_{DEP}{\nabla\left( {E \cdot E} \right)}}}} & (8)\end{matrix}$

where μ_(DEP) is the dielectrophoretic mobility of the particle and isdefined as:

$\begin{matrix}{\mu_{DEP} = {\frac{ɛ_{m}r^{2}}{3\eta}{{Re}\left( \frac{ɛ_{p}^{*} - ɛ_{m}^{*}}{{ɛ_{p}^{*} + {2ɛ_{m}^{*}}}\;} \right)}}} & (9)\end{matrix}$

Methods

Microfabrication Process

Deep Reactive Ion Etching (DRIE)

A silicon master stamp was fabricated on a <100> silicon substrate. AZ9260 (AZ Electronic Materials) photoresist was spun onto a clean siliconwafer and softbaked at 114 C for 45 seconds (FIG. 21 a). The wafer wasthen exposed to UV light for 45 seconds with an intensity of 12 W/mthrough a chrome plated glass mask. The exposed photoresist was thenremoved using Potassium based buffered developer AZ400K followed byanother hard baking at 115 C for 45 seconds (FIG. 21 b). Deep ReactiveIon Etching (DRIE) was used to etch the silicon master stamp to depthsranging from 50-100 microns (FIG. 21 c). The silicon master stamp wasthen cleaned with acetone to remove any remaining photoresist (FIG. 21d). The scalloping effect, a typical effect of the DRIE etching method,creates a surface roughness which is detrimental to the stampingprocess. In order to reduce the surface roughness, silicon oxide wasgrown on the silicon master using thermal oxidation and then was removed(FIG. 21 g-i).

PDMS

The liquid phase PDMS was made by mixing the PDMS monomers and thecuring agent in a 10:1 ratio (Sylgrad 184, Dow Corning, USA). Thebubbles in the liquid PDMS were removed by exposing the mixture tovacuum for an hour. A enclosure was created around the wafer usingaluminum foil in order to contain the PDMS on the wafer as well as toensure the proper depth for the PDMS portion of the device. The cleanPDMS liquid was then poured onto the silicon master and 15 minutes wasallowed for degassing. The PDMS was then cured for 45 min at 100 C (FIG.21 e) and then removed from the mold. Finally, fluidic connections tothe channels were punched with 15 gauge blunt needles (Howard ElectronicInstruments, USA).

Bonding

Microscope glass slides (3″×2″×1.2 mm, Fisher Scientific, USA) werecleaned with soap and water and rinsed with distilled water andisopropyl alcohol then dried with a nitrogen gun. The PDMS replica wasbonded with the clean glass slides after treating with oxygen plasma for40 s at 50 W RF power (FIG. 21 f). A schematic with dimensions andequivalent circuit model of the device is presented in FIG. 22 a. Theside channels are separated from the sample channel with 20 μm PDMSbarriers.

Experimental Setup

Pipette tips, inserted in the punched holes in the PDMS portion of thedevice, were used as reservoirs for fluidic connections to the channels.Pressure driven flow (10 to 15 μl/hr was provided by an imbalance in theamount of the sample in these reservoirs of the main channel. Aninverted light microscope (Leica DMI 6000B, Leica Microsystems,Bannockburn, IL) equipped with a digital camera (Hamamatsu EM-CCD C9100,Hamamatsu Photonics K. K. Hamamatsu City, Shizuoka Pref., 430-8587,Japan) was used to monitor cells in the main channel. Microfluidicdevices were placed in a vacuum jar for at least half an hour beforerunning the experiments to reduce priming issues and then the side andmain microchannels were filled with PBS and DEP buffer respectively.

Cells and Buffer

The THP-1 human Leukemia monocytes, MCF-7 breast cancer cells, andMCF-10A breast cells were washed twice and resuspended in a prepared DEPbuffer (8.5% sucrose [wt/vol], 0.3% glucose [wt/vol], and 0.725%[vol/vol] RPMI)(Flanagan, Lu et al. 2008). The electrical conductivityof the buffer was measured with a Mettler Toledo SevenGo proconductivity meter (Mettler-Toledo, Inc., Columbus, Ohio) to ensure thatits conductivity was 1000 μS/cm. These cells were observed to bespherical while they are in suspension. The measured cell diameters ofwith the corresponding standard deviations (n=30) of these cell aregiven in Table 2 below.

Electronics

A commercially available two-transistor inverter circuit (BXA-12576, JKLComponents Corp., USA) was modified to provide a high-frequency andhigh-voltage AC signal for the device (FIG. 2 b). The circuit relies onthe oscillation created by the two-transistors and passive components tocreate an AC voltage on the primary side of a transformer. This voltageis then stepped-up by the transformer to give a high-output voltage onthe secondary side to which the microfluidic device was connected.

The resonant frequency at which the circuit operates is highly dependanton the load impedance connected to the secondary side of thetransformer. Two high-voltage power supplies were fabricated withresonant frequencies of 85 kHz and 126 kHz. A DC input voltage wasprovided by a programmable DC power supply (PSP-405, Instek AmericaCorp., USA) which allowed adjustment of the output voltage by varyingthe input voltage. This technique allowed the output voltage of thepower supplies to be varied from approximately 100 Vrms to 500 Vrms. Athree-resistor voltage divider network, with a total impedance of onemegaohm, was added to the output of the inverter circuit in order toprovide a scaled (100:1) output voltage to an oscilloscope(TDS-1002B,Tektronix, USA) which facilitated monitoring the frequencyand magnitude of the signal applied to the microfluidic device. Allcircuitry was housed in a plastic enclosure with proper high-voltagewarnings on its exterior and connections were made to the microfluidicdevice using high-voltage test leads.

Translational and Rotational Velocity Measurement

The average velocity of the THP-1, MCF-7 and MCF-10A cells were measuredin the microfluidic device along the centerline a-b in FIG. 23 frompoint 1 to point 4. Time-lapse videos were recorded of the cells motionbefore and after applying an ac electric field through the platinumelectrodes inserted in the side channels. These recorded videos thenwere converted to JPEG files using the Leica software, (Leica DMI 6000B,LAS AF 1.6.3 Leica Microsystems, Bannockburn, Ill.), in order to measurethe traveling time of the target cells, for a known specific distance inthe microchannel, before and after inducing the electric field in themain microfluidic channel. Results are summarized below.

Numerical Modeling

The microfluidic device was modeled numerically in Comsol multi-physics3.4 using AC/DC module (Comsol Inc., Burlington, Mass. USA). Sincedielectrophoresis depends on the gradient of the electric field, ∇E=∇(∇Ø), it is necessary to determine the electric field distributionwithin a channel geometry. This is done by solving for the potentialdistribution, φ using the Laplace equation, ∇²Ø=0. The boundaryconditions used are prescribed uniform potentials at the inlet or outletof the side channels, and a zero derivative normal to the channel walls,∇ω·n=0, where n is the local unit vector normal to the walls.

The values for the electrical conductivity and permittivity of the PDMS,PBS, and DEP buffer that was used in this numerical modeling are givenin Table 1. PBS and DEP buffer electrical properties are used for theside and main microfluidic channels, respectively.

TABLE 1 Electrical properties of the materials and fluids. ElectricalProperties Electrical Conductivity Relative Electrical Materials (S/m)Permittivity PDMS 0.83 × e−12 2.65 PBS 1.4 80 DEP Buffer 0.01 80

The effect of the external voltage and the frequency on the gradient ofthe induced electric field has been studied. The gradient of theelectric field along the center line of the main channel is investigatednumerically for different applied voltages (100, 200, 350, and 500V) at85 kHz and for different frequencies (40, 85, 125, and 200 KHz) at 250Vrms applying voltage. Based on the available electronic circuit (250Vrms at 85 KHz), the electric field distribution and the gradient of theelectric field was mapped in the microfluidic device.

Results and Discussion

Numerical Results

FIG. 23 shows the surface and line plot of the gradient of the electricfield inside the main microfluidic channel at the intersection betweenthe main and the side channels. There is a high gradient of the electricfield at the corners (points 1 and 2) as well as point 3, which canprovide a strong DEP force. These results indicate that changes in thethickness of the PDMS barrier have a more significant effect on thegradient of the induced electric field inside the main channel thanchanges in the channel's geometry which is in agreement with theanalytical results.

In FIG. 23 b the gradient of the electric field along the line a-b isplotted for different applied frequencies (40, 85, 125, and 200 KHz) at250 Vrms. The effect of the total external voltage across themicrofluidic device on the gradient of the electric field (along theline a-b) is also investigated in FIG. 23 c. DEP response of the systemis plotted for four different voltages (100, 200, 350, and 500V) at 85kHz.

An increased gradient of the electric field can be obtained byincreasing the applied frequency or increasing the total applied voltagealthough it should be noted that adjusting the frequency will alsoaffect the Clasius-Mossotti factor of the microparticles and needs to beconsidered. Also the induced gradient of the electric field in the mainmicrofluidic channel is on the order of 10¹² (kg²·m·C⁻²·S⁻⁴) which isstrong enough for particle manipulations.

Based on this numerical modeling, the voltage drop across the 20 μm PDMSbarrier was 250V for an applied total voltage of 500V across themicrofluidic electrode channels. This voltage drop is lower than the400V break down voltage for a 20 μm PDMS channel wall. Thus, the DEPforce can be amplified by adjusting the input voltage with sometolerance.

Electric Field Surface Plot

FIGS. 24 a-c show the induced electric field intensity distributioninside the main microfluidic channel filled with the DEP buffer with aconductivity of 100 μS/cm. The highest electric field is induced at thezone of intersection between the main and the side channels and betweenthe PDMS barriers. FIG. 24 c also shows that with an applied AC electricfield of 250 Vrms and 85 kHz the electric field does not significantlyexceed 0.2 kV/cm in the main microfluidic channel.

Experimental Results

Cell Trapping-Contactless DEP Evidence

FIG. 25 shows the experimental results attained using MCF-7 breastcancer cells and THP-1 leukemia cells in the device. The behavior ofcells traveling through the device under static conditions was observedto be significantly different than when an electric field was applied tothe device. Three induced DEP responses were studied, rotation, velocitychanges, and chaining.

Under a pressure driven flow, without an applied electric field, it wasobserved that THP-1 leukemia and MCF-7 breast cancer cells flow throughthe main microfluidic channel from right to left without any disruptionor trapping. The cells were observed to be trapped, experiencing apositive DEP force, once an AC electric field at 85 KHz and 250 Vrms wasapplied. These results indicate that these cells have positiveClausius-Mossotti factor at 85 kHz frequency. Their velocity decreasedat the intersection between the main and the side channels where thethin PDMS barriers are located. With the same electrical boundaryconditions no trapping or cell movement disruption for MCF-10A normalbreast cells was observed. However, these cells were trapped once anelectric field at 125 kHz and 250 Vrms was applied.

Since the positive DEP force in the main microchannel depends on theelectrical properties of the cells, different cell lines experiencedifferent forces at the same electrical boundary conditions (externalvoltage and frequency) in the same buffer. Cell bursting or lysis wasnot observed during contactless DEP trapping.

Translational Velocity

The cells were observed to move faster along the centerline of thesample channel in FIG. 23 a from point 5 to point 1 when the electricfield was applied as compared to their velocity due to pressure drivenflow. As shown in FIG. 23, the magnitude of the DEP force is high atpoint 1. Because the DEP force is positive at 85 kHz, the cells areattracted to this point. Therefore, as the cells approach point 1 fromthe right, the positive DEP force is in the direction of the pressuredriven flow, causing the cells to move faster down the channel.Conversely, the average velocity of the cells in the area between thethin PDMS barriers (from 1 to 4) decreases when the voltage is appliedbecause the positive DEP force now acts in the opposite direction of thepressure driven flow.

Table 2 compares the induced velocities of the cells with respect totheir velocity under pressure driven flow. The normalized velocity(Uon/Uoff) for the three cell lines under the same electrical boundaryconditions (250 Vrms at 85 kHz) are also reported in FIG. 26. Theresults show that there is a statistically significant difference in thecells velocity when the field is applied. Furthermore, when theexperiments are normalized for comparison, the results suggest that thistechnique can be used to differentiate cells based on their electricalproperties.

TABLE 2 The measure average velocity from point 1 to point 4 (FIG. 23)of five different cells before and after applying the electric field atthe zone of trapping. Cell Velocity Diameter Uon Uoff Uoff-Uon Uon/ ΩCell line (μm) (μm/s) (μm/s) (μm/s) Uoff (rad/s) THP-1 15.4 ± 2   240 ±13 392 ± 21 152 ± 19  0.61  8.1 ± 0.66 MCF-7 18.5 ± 2.5 387 ± 7  476 ±17 89 ± 17 0.81 19.4 ± 2.9 MCF-10A 18.2 ± 2.1 310 ± 17 313 ± 16  3 ± 240.99 N.A.

The same experiments with the same buffers and electrical boundaryconditions were performed on MCF-10A breast cells without noticeabletrapping or disruption, which shows that the electrical properties ofthe normal breast cells are different compared to the MCF-7 breastcancer cells. It also shows the sensitivity of the contactless DEPtechnique to isolate cells with close electrical properties.

There was a great tendency for cells to move towards the corners in themain channel. This agrees with the numerical results, which show thereis a high gradient of the induced electric filed at the corners, whichcauses a strong positive DEP force and pulls cells towards these zonesof the main microfluidic channel.

Rotational Velocity

Cell rotation in the main channel at the zone of trapping and betweenthe thin PDMS barriers was present with an applied electric field. Therotational velocity of the cell is a function of its electricalproperties, the medium permittivity, the medium dynamic viscosity aswell as the properties of the electric field. The rotational velocity ofthe trapped THP-1, and MCF-7 cancer cells was measured in differentexperiments at one spot of the main microfluidic channel. No cellrotation was observed without an applied electric field. The reportedrotational velocities in Table 2 are the average rotational velocitiesof five different cells of each of the cancer lines. These results implythat the average rotation velocities of the THP-1 and MCF-7 cancer celllines are significantly different. Cell rotation for the MCF-10A cellswith the same electrical boundary conditions in the same buffer solutionwas not observed.

Pearl-Chain

Cell aggregation and chain formation in DEP experiments with an AC fieldhave been frequently observed and can be attributed to dipole-dipoleinteractions as well as local distortions of the electric field due tothe cells' presence (28, 29, 52, 62). Particles parallel to the electricfield attract each other because of this dipole-dipole force, resultingin pearl-chaining of the trapped cells in the direction of the electricfield in the microfluidic channel. The cell chain formation was observedfor the MCF-7 and THP-1 cancer cell lines in the experiments with anapplied AC electric filed at 85 KHz and 250 Vrms (FIG. 27).

Conclusion

This Example demonstrates a new technique for inducing electric fieldsin microfluidic channels in order to create a dielectrophoretic force.The method relies on the application of a high-frequency AC electricsignal to electrodes that are capacitively coupled to a microfluidicchannel. In the subject device, the geometry of the electrodes andchannels create the spatial non-uniformities in the electric fieldrequired for DEP. Three separate DEP responses were observed in thedevice, namely, translational velocity, rotational velocity, andchaining. In order to observe the devices effects in these threecategories, three different cell lines were inserted into the devicesand their individual responses recorded. Each cell line exhibited aresponse unique to its type due to the cell's specific electricalproperties. This result highlights the ability of this technique todifferentiate cells by their intrinsic electrical properties.

This technique may help overcome many of the challenges faced withtraditional iDEP and DEP. Because the induced electric field is not asintense as comparable methods and is focused just at the trapping zone,it is theorized that the Joule heating within the main microfluidicchannel is negligible. This could mitigate the stability and robustnessissues encountered with conventional iDEP (39), due the conductivitydistribution's strong dependence on temperature. Furthermore, challengesassociated with cell lysing due to high temperatures (37) orirreversible electroporation due to high field strengths (50, 65) areovercome with the new design approaches disclosed herein.

Example 2 Selective Isolation of Live/Dead Cells Using ContactlessDielectrophoresis (cDEP)

Introduction

Isolation and enrichment of cells/micro-particles from a biologicalsample is one of the first crucial processes in many biomedical andhomeland security applications (1). Water quality analysis to detectviable pathogenic bacterium (2-6) and the isolation of rare circulatingtumor cells (CTCs) for early cancer detection (7-19) are importantexamples of the applications of this process.

Dielectrophoresis (DEP) is the motion of a particle in a suspendingmedium due to the presence of a non-uniform electric field (28, 29). DEPutilizes the electrical properties of the cell/particle for separationand identification (29, 66). The physical and electrical properties ofthe cell, the conductivity and permittivity of the media, as well as thegradient of the electric field and its applied frequency are substantialparameters determining a cell's DEP response.

One unique advantage of DEP over existing methods for cell separation isthat the DEP force is strongly dependent on cell viability. The cellmembrane, which is normally impermeable and highly insulating, typicallybecomes permeable after cell death (31). This results in the release ofions from the cytoplasm through the structural defects in the dead cellmembrane and the cell conductivity will increase dramatically (32). Thisalteration in electrical properties after cell death make DEP live/deadcell separation and isolation possible.

The utilization of DEP to manipulate live and dead cells has previouslybeen demonstrated through several approaches. To start, Suehiro et al.were able to utilize dielectrophoretic impedance measurements toselectively detect viable bacteria (67). Conventional interdigitatedelectrode DEP micro devices have also been used to separate live andheat-treated Listeria cells (68). Huang et al. investigated thedifference in the AC electrodynamics of viable and non-viable yeastcells through DEP and electrorotation experiments (69) and a DEP-basedmicrofluidic device for the selective retention of viable cells inculture media with high conductivity was proposed by Docoslis et al.(70).

Insulator-based dielectrophoresis (iDEP) has also been employed toconcentrate and separate live and dead bacteria for water analysis (2).In this method, electrodes inserted into a microfluidic channel createan electric field which is distorted by the presence of insulatingstructures. The devices can be manufactured using simple fabricationtechniques and can be mass-produced inexpensively through injectionmolding or hot embossing (35, 36). iDEP provides an excellent solutionto the complex fabrication required by traditional DEP devices however,it is difficult to utilize for biological fluids which are highlyconductivity. The challenges that arise include joule heating and bubbleformation (37). In order to mitigate these effects, oftentimes theelectrodes are placed in large reservoirs at the channel inlet andoutlet. Without an additional channel for the concentrated sample (36),this could re-dilute the sample after it has passed through aconcentration region.

The development a robust, simple, and inexpensive technique to performDEP, termed “contactless dielectrophoresis” (cDEP) is described herein.This technique provides the non-uniform electric fields in microfluidicchannels required for DEP cell manipulation without direct contactbetween the electrodes and the sample (40). In this method, an electricfield is created in the sample microchannel using electrodes insertedinto two conductive microchambers, which are separated from the samplechannel by thin insulating barriers. These insulating barriers exhibit acapacitive behavior and therefore an electric field can be produced inthe main channel by applying an AC field across the barriers (40).

The absence of contact between the electrodes and the sample fluidprevents problems associated with more conventional approaches to DEPand iDEP including contamination, electrochemical effects, bubbleformation, and the detrimental effects of joule heating (33). Similar toiDEP, cDEP lends itself to a much simpler fabrication procedure. Devicesare typically molded from a reusable silicon master stamp that has beenfabricated from a single mask lithographic process. Once the masterstamp has been fabricated, cDEP devices can be produced from the stampoutside of the cleanroom environment, allowing for rapid, massfabrication of cDEP microfluidic devices.

As is shown below, the abilities of cDEP to selectively isolate andenrich a cell population was investigated. This was demonstrated throughthe separation of viable cells from a heterogeneous population alsocontaining dead cells. Two cDEP microfluidic devices were designed andfabricated out of polydemethilsiloxane (PDMS) and glass using standardphotolitography. The DEP response of the cells was investigated undervarious electrical experimental conditions in the range of the powersupply limitations. Human leukemia THP-1 viable cells were successfullyisolated from dead (heat treated) cells without lysing.

The separation of viable and nonviable cells is a critical startingpoint for this new technology to move towards more advancedapplications. Optimization of these devices would allow for selectiveseparation of cells from biological fluids for purposes such as: thediagnosis of early stages of diseases, drug screening, samplepreparation for downstream analysis, enrichment of tumor cells toevaluate tumor lineage via PCR, as well as treatment planning (41-46).By using viable/nonviable separation as a model for these applications,a new generation of cDEP devices can be tailored around the resultsreported in this study.

Theory

The 3D schematic of the experimental set up and device 1 is shown inFIG. 28. The dominant forces acting on the cell/particle in themicrofluidic devices are shown in FIGS. 29( a) and 30(a). For particleslarger than 1 μm, the Brownian motion is negligible compared to the DEPforce1. The DEP force acting on a spherical particle can be described bythe following (1, 28, 71)

F _(DEP)=2πε_(m) r ³Re{K(ω)}∇(E _(rms) ·E _(rms))   (1)

where ε_(m) is the permittivity of the suspending medium, r is theradius of the particle, E_(rms) is the root mean square electric field.Re{K(ω)} is the real part of the Clausius-Mossotti factor K(ω) . TheClausius-Mossotti is given by

$\begin{matrix}{{K(\omega)} = \frac{ɛ_{p}^{*} - ɛ_{m}^{*}}{{ɛ_{p}^{*} + {2ɛ_{m}^{*}}}\;}} & (2)\end{matrix}$

where ε*_(p) and ε*_(m) are the particle and the medium complexpermittivity respectively. The complex permittivity is defined asfollows:

$\begin{matrix}{ɛ^{*} = {ɛ + \frac{\sigma}{j\omega}}} & (3)\end{matrix}$

where ε is the permittivity, σ a is the conductivity, j²=−1, and ω isthe angular frequency.

Using the complex permittivity given in equation (3) of the particle andmedium, the real part of Clausius-Mossotti factor is calculated asfollows (72):

${{Re}\left\lbrack f_{cm} \right\rbrack} = {\frac{\left( {\sigma_{p} - \sigma_{m}} \right)}{\left( {1 + {\omega^{2}\tau_{MW}^{2}}} \right)\left( {\sigma_{p} + {2\sigma_{m}}} \right)} + \frac{\omega^{2}{\tau_{MW}^{2}\left( {\varepsilon_{p} - \varepsilon_{m}} \right)}}{\left( {\varepsilon_{p} - {2\varepsilon_{m}}} \right)}}$

For cells, the complex permittivity can be estimated using a singleshell model, which is given by

$\begin{matrix}{ɛ_{p} = {ɛ_{mem}\frac{\gamma^{3} + {2\left( \frac{ɛ_{i} - ɛ_{mem}}{ɛ_{i} + {2ɛ_{mem}}} \right)}}{\gamma^{3} - \left( \frac{ɛ_{i} - ɛ_{mem}}{ɛ_{i} + {2ɛ_{mem}}} \right)}}} & (5)\end{matrix}$

where γ=r/(r−d), r is the particle radius, d is the cell membranethickness, ε_(i) and ε_(mem) are the complex permittivites of thecytoplasm and the membrane, respectively (1, 72).

The parabolic velocity profile in the microchannel, shown in FIGS. 2( a)and 3(a), is due to the low Reynolds number pressure driven flow acrossthe main channel. Assuming the cell as a spherical particle, thehydrodynamic drag force due to cell translation is given by

f _(Drag)=6ηrπ(u _(p) −u _(f))   (6)

where r is the particle radius, η is the medium viscosity, u_(p) is thevelocity of the particle, and u_(f) is the medium velocity.

Others have shown that for micro particles moving in viscousenviroments, the inertial forces are negligible (73). The characteristictime for a spherical particle suspended in fluid is reported to be(2ρr²)/(9η) , where ρ is the density of the medium, r is radius of theparticle, and η is the viscosity of the medium.

For THP-1 cells with with 15.4±2 μm diameter (40) this characteristictime would be 12 μs, which is orders of magnitude smaller than the timescale of the external forces and the experimental observations. Thevelocity of the particle is determined by a balance between the DEPforce and Stoke's drag force. The relationship is given by

u _(p) =u _(f)−μ_(DEP)∇(E·E)   (7)

where μ_(DEP) is the dielectrophoretic mobility of the particle and isdefined as:

$\begin{matrix}{\mu_{DEP} = {\frac{ɛ_{m}r^{2}}{3\eta}{{Re}\left\lbrack f_{CM} \right\rbrack}}} & (8)\end{matrix}$

Methods

Fabrication

A silicon master stamp was fabricated on a <100> silicon substratefollowing the previously described process 32. Deep Reactive Ion Etching(DRIE) was used to etch the silicon master stamp to a depth of 50 μm.Silicon oxide was grown on the silicon master using thermal oxidationfor four hours at 1000° C. and removed with HF solvent to reduce surfacescalloping. Liquid phase polydimethylsiloxane (PDMS) was made by mixingthe PDMS monomers and the curing agent in a 10:1 ratio (Sylgrad 184, DowCorning, USA). The degassed PDMS liquid was poured onto the siliconmaster, cured for 45 min at 100° C., and then removed from the mold.Fluidic connections to the channels were punched using hole punchers(Harris Uni-Core, Ted Pella Inc., Redding, Calif.; 1.5 mm for the sidechannels and 2.0 mm for the main channel inlet and outlet. Microscopeglass slides (75 mm×75 mm×1.2 mm, Alexis Scientific) were cleaned withsoap and water, rinsed with distilled water, ethanol, isopropyl alcohol,and then dried with compressed air. The PDMS mold was bonded to cleanglass after treating with air plasma for 2 minutes. Schematics of thedevices with dimensions are shown in FIGS. 29( a) and 30(a).

Cell Preparation

The live samples of THP-1 human leukemia monocytes were washed twice andresuspended in a buffer used for DEP experiments (8.5% sucrose [wt/vol],0.3% glucose [wt/vol], and 0.725% [wt/vol] RPMI 43) to 106 cells/mL. Thecell samples to be killed were first pipetted into a conical tube andheated in a 60° C. water bath for twelve minutes; an adequate timedetermined to kill a majority of the cell sample.

To enable simultaneous observation under fluorescent microscope, cellswere stained using a LIVE/DEAD® Viability/Cytotoxicity Kit for mammaliancells (Molecular Probes Inc.). Calcein AM, which is enzymaticallyconverted to green fluorescent calcein, was added to the live cellsample at 2 μL per ml of cell suspension. Ethidium homodimer-1 (EthD-1)was added to the dead cell sample at 6 μL per ml of cell suspension.This can only pass through damaged cell membranes and upon nucleicacid-binding produces a red fluorescence.

The two samples were then vortexed for 5 minutes, washed once andresuspended in DEP buffer. The live and dead suspensions were then mixedtogether in one conical tube with a final concentration of 106 cells/mLand final conductivity of 110-115 μm measured with a SevenGo Proconductivity meter (Mettler-Toledo, Inc., Columbus, Ohio). Live and deadcells were indistingushable under bright field evaulation.

Experimental Set-Up

The microfluidic devices were placed in a vacuum jar for 30 minutesprior to experiments to reduce problems associated with priming. Pipettetips were used to fill the side channels with Phosphate Buffered Saline(PBS) and acted as reservoirs. Aluminum electrodes were placed in theside channel reservoirs. The electrodes inserted in side channels 1 and2 of device 1 (FIG. 29 a) were used for excitation while the electrodesinserted in side channels 3 and 4 were grounded. The electrodes insertedin side channel 1 of device 2 (FIG. 30 a) were used for excitation whilethe electrodes inserted in side channel 2 were grounded. Thin walledTeflon tubing (Cole-Parmer Instrument Co., Vernon Hills, Ill.) wasinserted into the inlet and outlet of the main channel. A 1 ml syringecontaining the cell suspension was fastened to a micro-syringe pump(Cole Parmer, Vernon Hills, Ill.) and connected to the inlet tubing.Once the main channel was primed with the cell suspension, the syringepump was set to 0.02 mL/hr; equivalent to a velocity of 556 μm/sec fordevice 1 and 222 μm/sec for device 2. This flow rate was maintained for5 minutes prior to experiments.

An inverted light microscope (Leica DMI 6000B, Leica Microsystems,Bannockburn, Ill.) equipped with color camera (Leica DFC420, LeicaMicrosystems, Bannockburn, Ill.) was used to monitor the cells flowingthrough the main channel. Once the flow rate of 0.02 ml/hr wasmaintained for 5 minutes an AC electric field was applied to theelectrodes.

Device 1: Experiments were conducted at 50 Vrms, 75 Vrms, 100 Vrms, 125Vrms and 150 Vrms. Trapping boundary conditions for this device weredetermined through visual inspection of the cells passing through themain channel. At each voltage, frequency was recorded for 80% trappingand the beginning of cell lyses. Significant lysing was considered to bewhen at least 10% of the cell population became lysed. The electricfield was maintained for 30 seconds during each experiment. Eight trialswere conducted at each voltage and corresponding frequencies wererecorded where 80% trapping was observed.

Device 2: Trapping efficiency for this device was determined forvoltages of 20 Vrms, 30 Vrms, 40 Vrms, 50 Vrms and frequencies of 200kHz, 300 kHz, 400 kHz, 500 kHz at a constant flow rate of 0.02 mL/hr.Experimental parameters were tested at random to mitigate any variationin cell concentration, flow rate, device functionality and otherexperimental variables. Additionally, trapping efficiency was calculatedat 0.02 mL/hr, 0.04 mL/hr, 0.06 mL/hr, and 0.08 mL/hr, with electricalparameters held constant at 500 kHz and 30 Vrms. Electrical parameterswere selected randomly for each experiment for a total of five trials ateach combination. The electric field was maintained for 30 secondsduring each experiment. During the 30 second interval, all cellsentering the trapping region of the device (the region containingpillars in the main channel) were counted, representing the total numberof cells.

Electrical Equipment

AC electric fields were applied to the microfluidic devices using acombination of waveform generation and amplification equipment. Waveformgeneration was performed by a function generator (GFG-3015, GW Instek,Taipei, Taiwan) whose output was then fed to a wideband power amplifier(AL-50HF-A, Amp-Line Corp., Oakland Gardens, N.Y). The wideband poweramplifier performed the initial voltage amplification of the signal andprovided the necessary output current to drive a custom-woundhigh-voltage transformer (Amp-Line Corp., Oakland Gardens, N.Y). Thistransformer was placed inside a grounded cage and attached to thedevices using high-voltage wiring. Frequency and voltage measurementswere accomplished using an oscilloscope (TDS-1002B, Tektronics Inc.Beaverton, Oreg.) connected to a 100:1 voltage divider at the output ofthe transformer.

Numerical Modeling

The electric field distribution and its gradient ∇E=∇(∇Ø) were modelednumerically in Comsol multi-physics 3.5 using the AC/DC module (ComsolInc., Burlington, Mass., USA). This is done by solving for the potentialdistribution, Φ, using the governing equation, ∇·(σ*∇Ø)=0, where σ* isthe complex conductivity (σ*=σ+jωε) of the sub-domains in themicrofluidic devices. The boundary conditions used are prescribeduniform potentials at the inlet or outlet of the side channels.

The values for the electrical conductivity and permittivity of the PDMS,PBS, and DEP buffer that were used in this numerical modeling are givenin Table 3. PBS and DEP buffer electrical properties are used for theside and main microfluidic channels, respectively. The induced DEPeffect inside the main channel was investigated for a range offrequencies and voltages. The gradient of the electric field along thecenter line (y=0) of the main channel as well as y=50 μm and y=100 μmwas investigated numerically.

TABLE 3 Electrical properties of the materials and fluids. ElectricalProperties Electrical Conductivity Relative Electrical Materials (S/m)Permittivity PDMS 0.83 × 10⁻¹² 2.65 PBS 1.4 80 DEP Buffer 0.01 80

Results and Discussion

Device 1: The geometry of device 1 allowed for the rapid simulation ofDEP effects within the sample microchannel which could then be verifiedthrough an efficient fabrication and experimentation procedure. Thegradient of the electric field along the center line of the main channelof device 1 was numerically modeled and the results are plotted in FIG.29 b. FIG. 29 b also shows that the maximum gradient of the electricfield occurs at the terminations of the side channels. The dependance ofthe gradient of the electric field in the main channel on distance fromthe channel wall is shown in FIG. 29 c. These numerical results indicatethat the gradient of the electric field and thus the DEP effect isstrongly related to the channel geometry.

Conclusions drawn from the numerical modeling of device 1 were verifiedthrough direct experimentation. Live cell concentration and trapping wasobserved for the electrical boundary conditions that were previouslysimulated (V1=V2=50 Vrms at 220 kHz, 100 Vrms at 152 kHz, and 150 Vrmsat 142 kHz and V3=V4=Ground). A large DEP response was achieved with anapplied voltage of 150 Vrms at 142 kHz, minoring the numerical modelingshown in FIG. 29 b. The majority of cell trapping within the deviceoccurred at the edges of the electrodes as predicted by numericalresults found in FIG. 29 b.

When 80% trapping was observed, cells closest to the channel wall weretrapped while those closer to the center of the channel were not; aresult predicted by the numerical modeling presented in FIG. 29 c. Thesesimulations further indicated that at low frequencies (≦100 kHz) thegradient of the electric field inside the main channel would not besufficient for DEP cell manipulation and this was confirmed in theexperiments. The minimum frequency necessary to achieve an 80% trappingefficiency is given in FIG. 31 a as a function of applied voltage. Celllysing was observed for 75 Vrms, 100 Vrms, 125 Vrms, and 150 Vrms at 296kHz, 243 kHz, 197 kHz, and 173 kHz respectively. No lysing was observedat 50 Vrms within the frequency limits of the power supply. Theconcentration of live THP-1 cells using a 150 kHz voltage signal at 100Vrms in device 1 is shown in FIG. 32.

Device 2: Numerical modeling proven valid for device 1 was used topredict the performance of device 2. The gradient of the electric fieldalong the x-axis (y=0) of the main channel of device 2 is plotted inFIG. 3 b. Again, for these electrical boundary conditions (V1=30 Vrms at200 kHz, 300 kHz, 400 kHz, and 500 kHz and V2=Ground) cell trapping wasobserved. Local maximums in the gradient of the electric field occurredin line with the edges of the insulating pillars while the minimumgradient was experienced as cells passed through the region between twopillars. The highest electric field gradient was observed to occur atthe two insulating pillars which had edges in the center of the device.The electric field gradients in the center of device 2 along the y-axis(x=0) are shown in FIG. 30 c and the highest gradient was observed inline with the edges of the insulating pillars. It should be noted thatthe maximum gradient is observed at y=+/−83.5 μm and cells travelingthrough the exact center of the device (along the x-axis) experience alower DEP force than those just off-center. The electric field gradientwithin the channel increased with applied signal frequency from 200 kHzto 500 kHz. This increase in gradient is not linear and these parametersrepresent the limitations of the current electrical setup.

Theoretically, device 2 has a maximum gradient of electric field withinthe channel occuring between 600 kHz and 700 kHz as seen in FIG. 31 d.Above this frequency, leakages in the system begin to dominate theresponse and the electric field within the channel drops off.

Live THP-1 cells were observed to experience positive DEP force at thereported frequencies and the DEP force applied on dead cells appeared tobe negligible. In device 2, the majority of cell trapping was observedin the region between the first two columns of insulating barriers at0.02 mL/hour. However, the distribution of trapped cells became moreuniform at higher flow rates. At 0.02 mL/hour, trapping efficienciesgreater than 90% were observed at all tested frequencies (200 kHz, 300kHz, 400 kHz, and 500 kHz). However, lysing was seen at all frequencieswhen a voltage of 50 Vrms was applied. At the highest two frequencies,lysing was seen at 40 Vrms and over 10% of the cells lysed at 50 Vrms(FIG. 31 c). Aside from lysing, the maximum voltage which may be appliedto these devices is determined by the electrical breakdown voltage ofthe PDMS composing the barriers. These results suggest that theperformance of the cDEP devices is comparable to and maybe able toexceed what is currently attainable and has been reported with DEP oriDEP 44-47.

In device 2, a maximum of 50 Vrms was applied to the inlets of theelectrode channels. In device 2, a maximum of 50 Vrms at 500 kHz signalwas applied to the inlets of the electrode channels. Because the samplechannel is non-uniform, it was found through the numerical results thatthe actual electric field experienced by cells within the channel wasbetween 20 V/cm and 200 V/cm. However, there are minute regions at thesharp corners inside the main channel with a high electric fieldintensity (350 V/cm) that induces electroporation (IRE), which is whatwas observed during the experiments. This was caused by the dramaticchange in the thickness of the PDMS barrier in those locations. It wasin these small regions which cell lysing was most commonly seen.

Trapping efficiency experiments for higher flow rates were conducted at500 kHz and 30 Vrms because these parameters yielded a high trappingefficicncy of 89.6% at 0.02 mL/hour. Trapping efficiency was reduced byan increase in flow rate and reached a minimum of 44.8% (+/−14.2) at 0.8mL/hour (FIG. 31 b). Flow rates greater than 0.1 mL/hour were notreported due to limitations of the recording software that resulted inthe inability to accurately count the number of cells entering andexiting the trapping region of the device.

Due to the capacitance effect of the PDMS barriers in cDEP devices, thecorresponding gradient of the electric field for voltage-frequency pairsare different for each design. These devices were designed to provide asufficient gradient of the electric field for DEP cell manipulationwithin the limitations of the power supply and the PDMS breakdownvoltage. The high trapping efficiency makes device 2 an optimal designfor selective entrapment and enrichment of cell samples. This process isdepicted in FIG. 33; initially live cells and dead cells passed throughthe trapping region due to pressure driven flow (FIG. 33). Live cellswere selectively concentrated in the trapping region under theapplication of a 500 kHz, 40 Vrms signal (FIG. 33 b). Under theseparameters, the DEP force on the dead cells was not sufficient toinfluence their motion and they passed through the trapping region. Theenriched sample of live cells can be controllably released for lateranalysis once the electric field is turned off (FIG. 33 c).

Conclusion

This work has demonstrated the ability of cDEP to selectivelyconcentrate specific cells from diverse populations through theseparation of viable cells from a sample containing both viable andnon-viable human leukemia cells. Repeatability, high efficiency,sterility, and an inexpensive fabrication process are benefits inherentto cDEP over more conventional methods of cell separation. This methodis also unique in that direct evaluation is possible with little or nosample preparation. The resulting time and material savings areinvaluable in homeland security and biomedical applications. GivencDEP's numerous advantages, the technique has tremendous potential forsample isolation and enrichment for drug screening, disease detectionand treatment, and other lab-on-a-chip applications.

Example 3 Biological Particle Enrichment Utilizing cDEP

Introduction

The selective separation of target particles from a sample solution isan indispensable step in many laboratory processes [1]. Sensitiveanalysis procedures, especially those in the biomedical field, oftenrequire a concentration procedure before any analysis is performed.Several methods to perform this concentration have arisen including:density gradient based centrifugation or filtration [57], fluorescentand magnetic activated cell sorting, cell surface markers [55], andlaser tweezers [79]. While, each of these techniques is unique in itsinherent advantages and disadvantages, all are forced to compromisebetween high sample throughput and highly specific isolation. The moreselective of these techniques oftentimes require extensive samplepreparation before being performed. If the automation of laboratoryanalysis procedures is to be facilitated, a concentration techniquecapable of high sample throughput as well as highly specificconcentration is critical.

Dielectrophoresis (DEP), or the motion of a particle due to itspolarization in a non-uniform electric field, has shown great potentialas a method for sample concentration [28, 29]. Typically, sampleconcentration through DEP involves the placement of an array ofinterdigitated electrodes under a microfluidic channel through which thesample fluid is passing. This electrode array creates a non-uniformelectric field in the channel with which passing cells ormicro-particles interact. DEP-based concentration techniques benefitfrom the fact that particles are isolated based upon their physicalcharacteristics; allowing these techniques to be extremely specificwithout extensive sample preparation.

Microdevices employing interdigitated electrode arrays have proven thetechnique to be a viable method to rapidly and reversibly isolate cellsand micro-particles from a solution. Examples of the successful use ofDEP include the separation of human leukemia cells from red blood cellsin an isotonic solution [7] and the entrapment of human breast cancercells from blood [8]. DEP has additionally been found effective toseparate neuroblastoma cells from HTB glioma cells [9], isolate cervicalcarcinoma cells [10], K562 human CML cells [11], and to separate liveyeast cells from dead [12].

Unfortunately, by requiring the fabrication of an electrode array withinthe microfluidic channel, traditional DEP does not lend itself to massfabrication techniques such as injection molding. Insulator-basedDielectrophoresis (iDEP) seeks to simplify the fabrication required toperform DEP-based concentration in order to facilitate more widespreadusage. iDEP relies upon the presence of insulating structures in themicrofluidic channel to create non-uniformities in the electric fieldnecessary for DEP [38, 51]. These insulating structures are typicallypatterned in the same process as the microfluidic channel itself; thus,iDEP naturally lends itself to mass production systems such as injectionmolding and hot embossing [35]. iDEP has been demonstrated incombination with other forms of on-chip analysis, such as impedancedetection [36], to form fully integrated systems.

While iDEP provided an excellent solution to the complex fabricationrequired by traditional DEP devices, it is difficult to utilize forbiological fluids. The high electric field intensity employed by iDEPproduces undesirable results such as joule heating, bubble formation,and electrochemical effects when the sample solution is of highconductivity [37]. In addition, the electrode placement at the channelinlet and outlet necessitates the presence of large reservoirs at theselocations to mitigate electrolysis effects. These reservoirs have thenegative consequence of re-diluting the sample after it has passedthrough the region of concentration, further complicating the extractionof a sample for off-chip analysis. For DEP to truly represent anattractive alternative to traditional sample concentration techniques,it must be devoid of these negative influences upon the sample and yetretain a simplified fabrication process.

A third manifestation of DEP, contactless dielectrophoresis (cDEP),employs the simplified fabrication processes of iDEP yet lacks theproblems associated with the electrode-sample contact [80]. cDEP reliesupon reservoirs filled with highly conductive fluid to act as electrodesand provide the necessary electric field. These reservoirs are placedadjacent to the main microfluidic channel and are separated from thesample by a thin barrier of a dielectric material as is shown in FIG. 1h. The application of a high-frequency electric field to the electrodereservoirs causes their capacitive coupling to the main channel and anelectric field is induced across the sample fluid. Similar totraditional DEP, cDEP exploits the varying geometry of the electrodes tocreate spatial non-uniformities in the electric field. However, byutilizing reservoirs filled with a highly conductive solution, ratherthan a separate thin film array, the electrode structures employed bycDEP can be fabricated in the same step as the rest of the device; hencethe process is conducive to mass production [80].

A cDEP device is presented that demonstrates the enrichment abilitiesand rapid fabrication advantages of the cDEP technique. A microfluidicdevice was fabricated by creating a PDMS mold of a silicon masterproduced by a single-mask photolithographic process. This device hasshown the ability of cDEP to separate live cells from dead [47] apowerful capability of DEP systems[67-70, 81]. In order to demonstratethe concentration abilities of cDEP, this microfluidic device was usedto enrich THP-1 human leukemia cells and 2-μm polystyrene beads from abackground media. The device exhibited the ability to concentrate THP-1cells through positive DEP and 2 μm beads via negative DEP. This is thefirst cDEP microfluidic device presenting negative DEP. Furthermore, theuse of a silicon master stamp allows for the large-scale reproduction ofthe device. These experiments illustrate that the use of cDEP as anexpedited process for sample concentration and enrichment, which mayhave an immense impact in biomedical and homeland security applicationswhere rapid, accurate results are extremely valuable.

Theory

The time-average dielectrophoretic force acting on a spherical particleexposed to a non-uniform electric field is described as [1, 28, 29, 71]

F _(DEP)=πε_(m) r ³Re[f _(CM) ]∇∥E∥ ²   (1)

where ε_(m) is the permittivity of the suspending medium, r is theradius of the particle, ∇∥E∥² defines the local electric field gradient,Re[ ] represents the real part, and f_(CM) is the Clausius-Mossottifactor given by

$\begin{matrix}{f_{CM} = \frac{ɛ_{p}^{*} - ɛ_{m}^{*}}{{ɛ_{p}^{*} + {2ɛ_{m}^{*}}}\;}} & (2)\end{matrix}$

where ε*_(p) and ε*_(m) are the particle and the medium complexpermittivity respectively.

The complex permittivity is defined as follows:

$\begin{matrix}{ɛ^{*} = {ɛ - {j\frac{\sigma}{\omega}}}} & (3)\end{matrix}$

where ε is the permittivity, a is the conductivity, j²=−1, and ω is theangular frequency. The hydrodynamic drag force on a spherical particledue to its translational movement in a suspension is given by:

f _(Drag)=6ηrπ(u _(p) −u _(f))   (4)

where r is the particle radius, η is the medium viscosity, u_(p) is thevelocity of the particle, u_(r) is the medium velocity. Assuming thatthe acceleration term can be neglected, the magnitude of the velocity ofthe particle is determined by a balance between the DEP force andStoke's drag force.

u _(p) =u _(f)−μ_(DEP)∇(E·E)   (5)

The above equations are valid for spherical micro-particles, however,others have demonstrated that similar equations can be attained forother geometries, e.g., cylindrical particles [82]. In addition,researchers have employed elegant shell models to determine aneffective/equivalent complex conductivity for a particle consisting ofseveral layers, e.g., a cell [83, 84].

The DEP force on a particle may be positive or negative depending on therelationship of the applied frequency to the particles DEP crossoverfrequency. DEP crossover frequency is the frequency in which the realpart of the Clausius-Mossotti (C.M.) factor is equal to zero and isgiven by [1, 72]

$\begin{matrix}{\omega_{c} = {\frac{1}{2\pi}\sqrt{\frac{\left( {\sigma_{m} - \sigma_{p}} \right)\left( {\sigma_{p} + {2\sigma_{m}}} \right)}{\left( {\varepsilon_{m} - \varepsilon_{p}} \right)\left( {\varepsilon_{p} + {2\varepsilon_{m}}} \right)}}}} & (6)\end{matrix}$

where ω_(c) is the crossover frequency and ρ_(p) and ρ_(m) are theconductivity of the particle and medium, respectively. This shows thatDEP can be used to differentiate micro-particles based on theirdifference in C.M. factor by adjusting the frequency.

Methods

Microfabrication

Deep Reactive Ion Etching (DRIE) was used to etch a <100> silicon waferto a depth of 50 μm (FIG. 34 a-d) to form the master stamp. Oxide wasthen grown on the silicon master using thermal oxidation and removedusing HF solvent to reduce surface “scalloping” caused by the DRIEprocess. This variation in the surface can greatly inhibit the removalof the cured mold from the stamp.

Liquid polydimethylsiloxane (PDMS) used for the molding process wascomposed of PDMS monomers and a curing agent in a 10:1 ratio (Sylgrad184, Dow Corning, USA). The mixture was de-gassed in a vacuum for 15minutes. The de-gassed PDMS liquid was then poured onto the siliconmaster and cured for 45 min at 100° C. (FIG. 1 e). The solidified PDMSwas removed from the mold and fluidic connections to the channels werepunched with 15 gauge blunt needles (Howard Electronic Instruments,USA). Cleaned glass microscope slides and the PDMS replica were bondedafter exposure to oxygen plasma for 40 s at 50 W RF power (FIG. 34 f). ASEM image of the trapping zone of the device replica on the siliconmaster is shown in FIG. 34 g. FIG. 1 h shows the fabricated device atthe zone of trapping. The main and electrode channels were filled withyellow and blue dyes respectively to improve imaging of the fluidicstructures. A schematic with dimensions is presented in FIG. 35. Thethickness of the PDMS barrier between the side channels and the mainchannel is 20 μm.

Cells/Beads and Buffer

Live samples of THP-1 human Leukemia monocytes were washed twice andresuspended in the prepared buffer (8.5% sucrose [wt/vol], 0.3% glucose[wt/vol], and 0.725% [wt/vol] RPMI) [74] to achieve 10⁶ cells/ml cellconcentration. The electrical conductivity of the buffer was measuredwith a Mettler Toledo SevenGo pro conductivity meter (Mettler-Toledo,Inc., Columbus, Ohio) to ensure that its conductivity was 1300 μS/cm.These cells were observed to be spherical with a diameter of ˜13 μm whenin suspension.

Carboxylate-modified polystyrene microspheres (Molecular Probes, Eugene,Oreg.) having a density of 1.05 mg/mm³ and diameters of 2 μm and 10 μmwere utilized at a dilution of 2:1000 from a 2% by wt. stock suspension.Bead suspensions were sonicated between steps of serial dilution andbefore use. The background solution was deionized water with aconductivity of 86 μS/cm.

Live THP-1 cells were stained using cell trace calcein red-orange dye(Invitrogen, Eugene, Oreg., USA). The stained cell sample and the 10 μmbeads sample were mixed in a ratio of 1:1.

Experimental Set-Up

The microfluidic devices were placed in a vacuum jar for 30 minutesprior to experiments to reduce problems associated with priming. Pipettetips inserted in the punched holes were used as reservoirs to fill theside channels with PBS. Pressure driven flow was provided in the mainchannel using a microsyringe pump. Inlet holes punched along the mainchannel of the device were connected to syringes via Teflon tubing(Cole-Parmer Instrument Co., Vernon Hills, Ill.). Once the main channelwas primed with the cell suspension, the syringe pump was set to 1 ml/hrsteadily decreasing the flow rate down to 0.02 ml/hr (20 μL/hr)equivalent to a velocity of ˜550 μm/sec. This flow rate was maintainedfor 1 minute prior to experiments. An inverted light microscope equippedwith color camera (DFC420, Leica DMI 6000B, Leica Microsystems,Bannockburn, Ill.) was used to monitor the cells flowing through themain channel. High-frequency electric fields were provided by awideband, high-power amplifier and transformer combination (Amp-LineCorp., Oakland Gardens, N.Y.) and signal generation was accomplishedusing a function generator (GFG-3015, GW Instek, Taipei, Taiwan).

Numerical Modeling

The electric field distribution and its gradient ∇E=∇(∇Ø) were modelednumerically in Comsol multi-physics 3.5 using the AC/DC module (ComsolInc., Burlington, Mass. USA). This is done by solving for the potentialdistribution, φ, using the Laplace equation, ∇·(σ*∇Ø)=0, where σ* is thecomplex conductivity of the sub-domains of the microfluidic device. Theboundary conditions used were prescribed uniform potentials at the inletor outlet of the side channels. The electrical conductivity and therelative electrical permittivity of PDMS have been reported as0.83×10⁻¹² S/m and 2.65 respectively (Sylgrad 184, Dow Corning, USA).The electrical conductivity of PBS and the DEP buffer are 1.4 S/m and1300 μS/cm respectively and a relative permittivity of 80.

Results

Numerical modeling was used to determine relevant experimentalconditions such as applied voltage and frequency. Experimental valuesfor the voltage and frequency must be chosen to provide sufficient DEPforce on the target particles without exceeding the dielectric breakdownvoltage of the PDMS barriers (280V for a 20 μm barrier). Due to thecapacitive properties of the thin PDMS barrier between the side channelsand the main channel, the induced electric field inside the main channelis strongly dependent on the frequency and the applied voltage. Hence, aminimum frequency is required to provide strong gradient of the electricfield with respect to a specific voltage for micro-particlemanipulation. A 70 V_(rms) sinusoid at 300 kHz was found to providesignificant DEP force in the microfluidic channel without damaging thedevice. This excitation signal was applied to the top two electrodes(electrodes 1 and 2) and the bottom two electrodes were grounded(electrodes 3 and 4). The electric field intensity surface plot in themain channel of the device at the experimental parameters is shown inFIG. 36 a. It is important to note that the electric field intensity didnot reach 0.1 kV/cm, the necessary field strength to kill cells throughirreversible electroporation. Electroporation is a phenomenon thatincreases the permeabilization of the cell membrane by exposing the cellto an electric field [85-87]. In irreversible electroporation, permanentpores open in the cell membrane which leads to cell death [86, 88].

The trapping regions and cell's trajectory through the microfluidicdevice can be predicted using the numerical modeling as DEP cellmanipulation is strongly dependent on the gradient of the electricfield. The highest gradient of the electric field is estimated to appearat the edges of the side channels as shown by numerical results found inFIG. 3 b. However, there is still a sufficient gradient of the electricfield at the middle of the channel to manipulate the micro-particles. Toclarify this, the same numerical results for the gradient of theelectric field surface plot, but with a different representing rangewere shown in FIG. 36 c.

The DEP force is acting on the cell/micro-particle in both x and ydirections. The gradient of the x-component of the electric field, whichcauses DEP force in the x-direction, is shown in FIG. 4 a for an appliedsignal of 70 V_(rms) and 300 kHz at three different distances from thechannel wall. In order to trap target cells, the x-component of the DEPforce should overcome the hydrodynamic drag force. The x-component ofthe DEP force along the centerline of the main channel is negligiblecompared to the DEP force along the channel wall. Furthermore, thisforce is the strongest along the edges of the side channel walls(x=−350, −150, 150, 350 μm). The y-component of the gradient of theelectric field at different distances from the origin (FIG. 35, x=0,150, 250, 350, and 450 μm) is also shown in FIG. 37 b. These resultsshow that the y-component of the DEP force is negligible for theparticles along the lines x=0 and x=250 compared to the other positionsand also indicate that y-component of the DEP force is the strongestalong the edges of the side channels (x=−350, −150, 150, 350 μm). Whilethe x-component of the DEP force along the centerline of the mainchannel is almost negligible (FIG. 37 a), the y-component of the DEPforce, will pull particles off the centerline of the main channel andtowards the channel walls in the case of positive DEP.

The effect of varying the electrode configuration on the gradient of theelectric field along the centerline of the main channel was alsoinvestigated. Four different configurations with the same appliedvoltage and frequency were studied and the results shown in FIG. 38. TheDEP effects caused by having electrodes 1 and 2 charged and electrodes 3and 4 grounded (case 1) are similar to the configuration with electrodes1 and 4 charged and electrodes 2 and 3 grounded (case 3). The same canbe said for the cases with electrodes 1, 2, and 4 charged (case 2) andelectrode 3 grounded or electrode 1 charged and electrode 2 grounded(case 4). The surface plot of the gradient of the electric field withrespect to these four cases of the electrode configurations were shownin FIG. 38 b.

These numerical results indicate that the electrode configuration has asubstantial effect on the gradient of the electric field and theresulting DEP cell manipulation. A benefit of this analysis is that onemay change the cell/particle manipulation strategy by changing theelectrode configurations. For example, the configuration used in case 4(electrodes on just one side of the main channel) can deflect the targetcell/particle trajectory in the main channel such that it leads to aspecific reservoir.

The validity of the numerical modeling was confirmed by demonstratingthe system's ability to concentrate particles through both positive andnegative DEP. Live THP-1 cells were observed to be trapped efficientlydue to positive DEP force at V₁=V₂=70 V_(rms) at 300 kHz, V₃=V₄=Ground(FIG. 39). Particles parallel to the electric field attract each otherdue to dipole-dipole interaction, resulting in pearl-chain formations ofthe trapped cells in the direction of the electric field [29, 53, 66].Referring to FIG. 36 b, particles concentrated through positive DEPshould show a predisposition to group at locations with a high gradientof the electric field, in this case at the edges of the electrodereservoirs. As can be seen in FIG. 39, this is indeed the case. Thepearl chain formations attach to the side wall at locations with a highgradient of the electric field and then spread towards the center of thechannel.

The selectivity of the device to differentiate two different particleswith almost the same size was also examined via separation of THP-1cells from 10 μm beads. The THP-1 cells were observed to be trapped at70 Vrms and 300 kHz and the 10 μm beads went through the main channelwithout significant DEP disturbance (FIG. 40). However, in order toincrease the trapping efficiency, the voltage and/or frequency of theapplied signal should be increased such that the particles passingthrough the middle of the channel experience strong DEP effect. At thesehigher voltage/frequencies that both cells and beads close to thechannel walls were observed to be trapped, reducing the device'sselectivity. This effect may be attributed to the non-uniform gradientof the electric field across the main channel and between the sidechannels.

Particle concentration through negative DEP was displayed using 2 μmbeads suspended in DI water at V1=V2=190 V_(rms) at 300 kHz andV3=V4=Ground. These experimental results are shown in FIG. 41. As isconsistent with a negative DEP response, the beads grouped in regionsaway from high gradients of the electric fields which, in this case, isin the centerline of the channel (FIGS. 40 b and c). The inability tofocus the microscope on all of the trapped beads simultaneouslyindicates that the beads were trapped at multiple heights in the mainchannel.

Discussion

The use of a straight channel in this design has several advantages overmore complicated configurations. The trajectory of a particle, withoutDEP influence, is easily predicted and the lack of detailed featuressimplifies production and replication of the devices. This same lack ofcomplicated features in the channel helps to mitigate fouling effectscaused by cell trapping. However, it should be noted that the DEP effectmay be reduced significantly at the middle of the channel for widerchannels. One method of addressing this negative effect is to useinsulating structures inside the main channel. These structures distortthe electric field and provide a sufficient gradient for DEPmanipulation of cells passing through the center of the channel. Thesetypes of designs may help increase the throughput and trappingefficiency of cDEP devices.

The device presented in this paper exhibited the concentration ofmicroparticles at specific trapping regions within the device during theapplication of an electric field. The removal of this electric fieldallows the trapped cells to flow from the device at an increasedconcentration and these cells may be diverted to a separate reservoiroff chip. This “trap and release” concentration strategy can also beincorporated with on-chip analysis systems by diverting the concentratedgroup of cells into a side channel as has been illustrated withiDEP[36].

Forthcoming generations of cDEP devices may also utilize a “chip andmanifold” configuration relying upon disposable, injection molded“chips” inserted into a reusable manifold containing the necessaryfluidic and electrical connections. This arrangement would allow metalelectrodes in the manifold to be re-used for thousands of experimentswhile shifting the manufacturing burden to the replication ofinexpensive fluidic chips. This use of polymer chips manufacturedthrough injection molding has been demonstrated previously for iDEP[36].

Conclusion

A microfluidic system was presented that illustrates the great potentialfor DEP-based concentration of biological particles without negativeeffects on the sample, extensive sample preparation, or complicatedfabrication procedures. Numerical modeling revealed the flexibility ofthis system's multiple electrode configurations to divert the particlesinto a desired trajectory and the device showed the ability toconcentrate micro-particles through both positive and negative DEP. Byrelying upon the particle's electrical properties to accommodateenrichment, cDEP should be able to achieve a high degree of specificitywithout extensive sample preparation.

The potential for batch fabrication illustrated in this work, combinedwith the high performance of the resulting devices makes cDEP anattractive candidate for pre-concentration processes in areas where bothrapid and highly accurate results of analyses are required.

Example 4 Continuous Separation of Beads and Human Red Blood Cells

The single layer device embodiment depicted in FIG. 42 consists of aT-channel 4217 and two electrode channels 4213, 4215. An exploded viewof the area in the box in FIG. 42A is shown in FIG. 42B. In theschematic in FIG. 42A, samples are introduced from left to right viapressure driven flow. When an AC electric signal of 100 Vrms at 400 kHzis applied across the fluid electrodes 4213, 4215, 4 micron beads can beisolated from 2 micron beads, concentrated, and released as shown inFIGS. 6 C and D. When an AC signal of 60V at 500 kHz is applied, humanred blood cells are separated from a buffer solution as shown in FIG.42E. In this device described in this example, particles arecontinuously separated from the bulk solution and diverted into aseparate microfluidic channel. Devices similar to this can be used toenhance microfluidic mixing.

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What is claimed is:
 1. A dielectrophoresis device comprising: a channelfor receiving a sample; a first electrode channel for receiving a firstelectrode; a first insulation barrier between the first electrodechannel and the channel for receiving a sample; a second electrodechannel for receiving a second electrode; and a second insulationbarrier between the second electrode channel and the channel forreceiving a sample.
 2. The dielectrophoresis device of claim 1, whereinthe channel for receiving the sample, the first electrode channel andthe second electrode channel are all formed in the same layer of thedevice.
 3. The dielectrophoresis device of claim 1, wherein the channelfor receiving a sample is linear.
 4. The dielectrophoresis device ofclaim 1, wherein the channel for receiving a sample is branched.
 5. Thedielectrophoresis device of claim 1, wherein the channel for receiving asample is T-shaped.
 6. The dielectrophoresis device of claim 1, whereinthe channel for receiving a sample, the first electrode channel and thesecond electrode channel are all formed in a single substrate layer, andwherein the first insulation barrier and second insulation barrier areformed by the substrate.
 7. The dielectrophoresis device of claim 1,wherein the first and second electrode channels are filled with aconductive solution selected from the group consisting of: phosphatebuffer saline, a conducting fluid, a conductive gel, a nanowire, aconductive paint, a polyelectrolyte, a conductive ink, a conductiveepoxy, and a conductive glues.
 8. The dielectrophoresis device of claim6, wherein the substrate layer is made from polydimethylsiloxane, glass,polyimide, polycarbonate, silicon or plastic.
 9. The dielectrophoresisdevice of claim 1, wherein an electrode channel or the channel forreceiving a sample comprises an insulation structure.
 10. Adielectrophoresis device comprising: a channel for receiving a sample ina first substrate layer; a first electrode channel for receiving anelectrode and a second electrode channel for receiving an electrode in asecond substrate layer; and an insulation barrier between the firstsubstrate layer and the second substrate layer.
 11. Thedielectrophoresis device of claim 10, wherein the first substrate layerand the second substrate layer are made from polydimethylsiloxane,glass, polyimide, polycarbonate, silicon or plastic.
 12. Thedielectrophoresis device of claim 10, wherein the insulation barrier ismade from polydimethylsiloxane, glass, polyimide, polycarbonate, siliconor plastic.
 13. The dielectrophoresis device of claim 10, wherein thechannel for receiving a sample is linear.
 14. The dielectrophoresisdevice of claim 10, wherein the channel for receiving a sample isbranched.
 15. The dielectrophoresis device of claim 10, wherein thechannel for receiving a sample is T-shaped.
 16. The dielectrophoresisdevice of claim 10, wherein the first and second electrode channels arefilled with a conductive solution selected from the group consisting of:phosphate buffer saline, a conducting fluid, a conductive gel, ananowire, a conductive paint, a polyelectrolyte, a conductive ink, aconductive epoxy, and a conductive glues.
 17. The dielectrophoresisdevice of claim 10, wherein an electrode channel or the channel forreceiving a sample comprises an insulation structure.
 18. Adielectrophoresis device comprising: a first electrode channel forconducting an electric current in a first substrate layer; a channel forreceiving a sample in a second substrate layer; a second electrodechannel for conducting an electric current in a third substrate layer; afirst insulation barrier between the first substrate layer and thesecond substrate layer; a second insulation barrier between the secondsubstrate layer and the third substrate layer.
 19. The dielectrophoresisdevice of claim 18, wherein the first and second substrate layers aremade from polydimethylsiloxane, glass, polyimide, polycarbonate, siliconor plastic.
 20. The dielectrophoresis device of claim 18, wherein theinsulation barrier is made from polydimethylsiloxane, glass, polyimide,polycarbonate, silicon or plastic.
 21. The dielectrophoresis device ofclaim 18, wherein the channel for receiving a sample is linear.
 22. Thedielectrophoresis device of claim 18, wherein the channel for receivinga sample is branched.
 23. The dielectrophoresis device of claim 18,wherein the channel for receiving a sample is T-shaped.
 24. Thedielectrophoresis device of claim 18, wherein the first and secondelectrode channels are filled with a conductive solution selected fromthe group consisting of: phosphate buffer saline, a conducting fluid, aconductive gel, a nanowire, a conductive paint, a polyelectrolyte, aconductive ink, a conductive epoxy, and a conductive glues.
 25. Thedielectrophoresis device of claim 18, wherein an electrode channel orthe channel for receiving a sample comprises an insulation structure.